Multi-phased, biodegradable and oesteointegrative composite scaffold for biological fixation of musculoskeletal soft tissue to bone

ABSTRACT

Methods and apparatuses are provided for musculoskeletal tissue engineering. For example, a scaffold apparatus is provided which comprises microspheres of selected sizes and/or composition. The microspheres are layered to have a gradient of microsphere sizes and/or compositions. The scaffold provides a functional interface between multiple tissue types.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application claims the benefit of U.S. Provisional Application Ser.No. 60/550,700, filed Mar. 5, 2004 and entitled “MULTI-PHASED,BIODEGRADABLE AND OSTEOINTEGRATIVE COMPOSITE SCAFFOLD FOR THE REPAIR OFMUSCULOSKELETAL TISSUE”, the entire contents of which are incorporatedherein by reference.

BACKGROUND

Throughout this application, certain publications are referenced. Fullcitations for these publications, as well as additional relatedreferences, may be found immediately preceding the claims. Thedisclosures of these publications are hereby incorporated by referenceinto this application in order to more fully describe the state of theart as of the date of the methods and apparatuses described and claimedherein.

This application relates to musculoskeletal tissue engineering. Forexample, a scaffold apparatus is discussed below which can serve as afunctional interface between multiple tissue types. Methods forpreparing a multi-phase scaffold are also discussed. Some exemplaryembodiments which include a soft tissue-bone interface are discussed.

As an example of a soft tissue-bone interface, the human anteriorcruciate ligament (ACL) is described below. The ACL and ACL-boneinterface are used in the following discussion as an example and to aidin understanding the description of the methods and apparatuses of thisapplication. This discussion, however, is not intended to, and shouldnot be construed to, limit the claims of this application.

The ACL consists of a band of regularly oriented, dense connectivetissue that spans the junction between the femur and tibia. Itparticipates in knee motion control and acts as a joint stabilizer,serving as the primary restraint to anterior tibial translation. Thenatural ACL-bone interface consists of three regions: ligament,fibrocartilage (non-mineralized and mineralized) and bone. The naturalligament to bone interface is arranged linearly from ligament tofibrocartilage and to bone. The transition results in varying cellular,chemical, and mechanical properties across the interface, and acts tominimize stress concentrations from soft tissue to bone.

The ACL is the most often injured ligament of the knee. Due to itsinherently poor healing potential and limited vascularization, ACLruptures do not heal effectively upon injury, and surgical interventionis typically needed to restore normal function to the knee.

Clinically, autogenous grafts based on either bone-patellar tendon-bone(BPTB) or hamstring-tendon (HST) grafts are often a preferred graftingsystem for ACL reconstruction, primarily due to a lack of alternativegrafting solutions. Current ACL grafts are limited by donor sitemorbidity, tendonitis and arthritis. Synthetic grafts may exhibit goodshort term results but encounter clinical failure in long-termfollow-ups, since they are unable to duplicate the mechanical strengthand structural properties of human ACL tissue. ACL tears and rupturesare currently commonly repaired using semitendinosus grafts. Althoughsemitendinosus autografts are superior, they often fail at the insertionsite between the graft and the bone tunnel. One of the major causes offailure in this type of reconstruction grafts is its inability toregenerate the soft-tissue to bone interface.

Despite their distinct advantages over synthetic substitutes, autogenousgrafts have a relatively high failure rate. A primary cause for the highfailure rate is the lack of consistent graft integration with thesubchondral bone within bone tunnels. The site of graft contact infemoral or tibial tunnels represents the weakest point mechanically inthe early post-operative healing period. Therefore, success of ACLreconstructive surgery depends heavily on the extent of graftintegration with bone.

ACL reconstruction based on autografts often results in loss offunctional strength from an initial implantation time, followed by agradual increase in strength that does not typically reach the originalmagnitude. Despite its clinical success, long term performance ofautogenous ligament substitutes is dependent on a variety of factors,including structural and material properties of the graft, initial grafttension, intrarticular position of the graft, as well as fixation of thegraft. These grafts typically do not achieve normal restoration of ACLmorphology and knee stability.

There is often a lack of graft integration with host tissue, inparticular at bony tunnels, which contributes to suboptimal clinicaloutcome of these grafts. The fixation sites at the tibial and femoraltunnels, instead of the isolated strength of the graft material, havebeen identified as mechanically weak points in the reconstructed ACL.Poor graft integration may lead to enlargement of the bone tunnels, andin turn may compromise the long term stability of the graft.

Increased emphasis has been placed on graft fixation, as post surgeryrehabilitation protocols require the immediate ability to exercise fullrange of motion, reestablish neuromuscular function and weight bearing.During ACL reconstruction, the bone-patellar tendon-bone orhamstring-tendon graft is fixed into the tibial and femoral tunnelsusing a variety of fixation techniques. Fixation devices include, forexample, staples, screw and washer, press fit EndoButton® devices, andinterference screws. In many instances, EndoButton® devices or Mitek®Anchor devices are utilized for fixation of femoral insertions. Staples,interference screws, or interference screws combined with washers can beused to fix the graft to the tibial region.

Recently, interference screws have emerged as a standard device forgraft fixation. The interference screw, about 9 mm in diameter and atleast 20 mm in length, is used routinely to secure tendon to bone andbone to bone in ligament reconstruction. Surgically, the knee is flexedand the screw is inserted from the para-patellar incision into thetibial socket, and the tibial screw is inserted just underneath thejoint surface. After tension is applied to the femoral graft and theknee is fully flexed, the femoral tunnel screw is inserted. Thisprocedure has been reported to result in stiffness and fixation strengthlevels which are adequate for daily activities and progressiverehabilitation programs.

While the use of interference screws have improved the fixation of ACLgrafts, mechanical considerations and biomaterial-related issuesassociated with existing screw systems have limited the long termfunctionality of the ligament substitutes. Screw-related laceration ofeither the ligament substitute or bone plug suture has been reported. Insome cases, tibial screw removal was necessary to reduce the painsuffered by the patient. Stress relaxation, distortion of magneticresonance imaging, and corrosion of metallic screws have providedmotivation for development of biodegradable screws based onpoly-α-hydroxy acids. While lower incidence of graft laceration wasreported for biodegradable screws, the highest interference fixationstrength of the grafts to bone is reported to be 475 N, which issignificantly lower than the attachment strength of ACL to bone. Whentendon-to-bone fixation with polylactic acid-based interference screwswas examined in a sheep model, intraligamentous failure was reported by6 weeks. In addition, fixation strength is dependent on quality of bone(mineral density) and bone compression.

Two insertion zones can be found in the ACL, one at the femoral end andanother located at the tibial attachment site. The ACL can attach tomineralized tissue through insertion of collagen fibrils, and thereexists a gradual transition from soft tissue to bone. The femoralattachment area in the human ACL was measured to be 113±27 mm² and136±33 mm² for the tibia insertion. With the exception of the mode ofcollagen insertion into the subchondral bone, the transition from ACL tobone is histologically similar for the femoral and tibial insertionsites.

The insertion site is comprised of four different zones: ligament,non-mineralized fibrocartilage, mineralized fibrocartilage, and bone.The first zone, which is the ligament proper, is composed of solitary,spindle-shaped fibroblasts aligned in rows, and embedded in parallelcollagen fibril bundles of 70-150 μm in diameter. Primarily type Icollagen makes up the extracellular matrix, and type III collagen, whichare small reticular fibers, are located between the collagen I fibrilbundles. The second zone, which is fibro-cartilaginous in nature, iscomposed of ovoid-shaped chondrocyte-like cells. The cells do not liesolitarily, but are aligned in rows of 3-15 cells per row. Collagenfibril bundles are not strictly parallel and much larger than thosefound in zone 1. Type II collagen is now found within the pericellularmatrix of the chondrocytes, with the matrix still made up predominantlyof type I collagen. This zone is primarily avascular, and the primarysulfated proteoglycan is aggrecan. The next zone is mineralizedfibrocartilage. In this zone, chondrocytes appear more circular andhypertrophic, surrounded by larger pericellular matrix distal from theACL. Type X collagen, a specific marker for hypertrophic chondrocytesand subsequent mineralization, is detected and found only within thiszone. The interface between mineralized fibrocartilage and subjacentbone is characterized by deep inter-digitations. Increasing number ofdeep inter-digitations is positively correlated to increased resistanceto shear and tensile forces during development of rabbit ligamentinsertions. The last zone is the subchondral bone and the cells presentare osteoblasts, osteocytes and osteoclasts. The predominant collagen istype I and fibrocartilage-specific markers such as type II collagen areno longer present.

For bone-patellar tendon-bone grafts, bone-to-bone integration with theaid of interference screws is the primary mechanism facilitating graftfixation. Several groups have examined the process of tendon-to-bonehealing.

Blickenstaff et al. (1997) evaluated the histological and biomechanicalchanges during the healing of a semitendinosus autograft for ACLreconstruction in a rabbit model. Graft integration occurred by theformation of an indirect tendon insertion to bone at 26 weeks. However,large differences in graft strength and stiffness remained between thenormal semi-tendinosus tendon and anterior cruciate ligament after 52weeks of implantation.

In a similar model, Grana et al. (1994) reported that graft integrationwithin the bone tunnel occurs by an intertwining of graft and connectivetissue and anchoring of connective tissue to bone by collagenous fibersand bone formation in the tunnels. The collagenous fibers have theappearance of Sharpey's fibers seen in an indirect tendon insertion.

Rodeo et al. (1993) examined tendon-to-bone healing in a canine model bytransplanting digital extensor tendon into a bone tunnel within theproximal tibial metaphysis. A layer of cellular, fibrous tissue wasfound between the tendon and bone, and this fibrous layer matured andreorganized during the healing process. As the tendon integrated withbone through Sharpey's fibers, the strength of the interface increasedbetween the second and the twelfth week after surgery. The progressiveincrease in strength was correlated with the degree of bone in growth,mineralization, and maturation of the healing tissue.

In most cases, tendon-to-bone healing with and without interferencefixation does not result in the complete re-establishment of the normaltransition, zones of the native ACL-bone insertions. This inability tofully reproduce these structurally and functionally different regions atthe junction between graft and bone is detrimental to the ability of thegraft to transmit mechanical stress across the graft proper and leads tosites of stress concentration at the junction between soft tissue andbone.

Zonal variations from soft to hard tissue at the interface facilitate agradual change in stiffness and can prevent build up of stressconcentrations at the attachment sites.

The insertion zone is dominated by non-mineralized and mineralizedfibrocartilage, which are tissues adept at transmitting compressiveloads. Mechanical factors may be responsible for the development andmaintenance of the fibrocartilagenous zone found at many of theinterfaces between soft tissue and bone. The fibrocartilage zone withits expected gradual increase in stiffness appears less prone tofailure.

Benjamin et al. (1991) suggested that the amount of calcified tissue inthe insertion may be positively correlated to the force transmittedacross the calcified zone.

Using simple histomorphometry techniques, Gao et al. determined that thethickness of the calcified fibrocartilage zone was 0.22±0.7 mm and thatthis was not statistically different from the tibial insertion zone.While the ligament proper is primarily subjected to tensile andtorsional loads, the load profile and stress distribution at theinsertion zone is more complex.

Matyas et al. (1995) combined histomorphometry with a finite elementmodel (FEM) to correlate tissue phenotype with stress state at themedial collateral ligament (MCL) femoral insertion zone. The FEM modelpredicted that when the MCL is under tension, the MCL midsubstance issubjected to tension and the highest principal compressive stress isfound at the interface between ligament and bone.

Calcium phosphates have been shown to modulate cell morphology,proliferation and differentiation. Calcium ions can serve as a substratefor Ca²⁺-binding proteins, and modulate the function of cytoskeletonproteins involved in cell shape maintenance.

Gregiore et al. (1987) examined human gingival fibroblasts andosteoblasts and reported that these cells underwent changes inmorphology, cellular activity, and proliferation as a function ofhydroxyapatite particle sizes. Culture distribution varied from ahomogenous confluent monolayer to dense, asymmetric, and multi-layers asparticle size varied from less than 5 μm to greater than 50 μm, andproliferation changes correlated with hydroxyapatite particles size.

Cheung et al. (1985) further observed that fibroblast mitosis isstimulated with various types of calcium-containing complexes in aconcentration-dependent fashion.

Chondrocytes are also dependent on both calcium and phosphates for theirfunction and matrix mineralization. Wuthier et al. (1993) reported thatmatrix vesicles in fibrocartilage consist of calcium-acidicphospholipids-phosphate complex, which are formed from actively acquiredcalcium ions and an elevated cytosolic phosphate concentration.

Phosphate ions have been reported to enhance matrix mineralizationwithout regulation of protein production or cell proliferation, likelybecause phosphate concentration is often the limiting step inmineralization. It has been demonstrated that human foreskin fibroblastswhen grown in micromass cultures and under the stimulation of lacticacid can dedifferentiate into chondrocytes and produce type II collagen.

Cheung et al. (1985) found a direct relationship betweenβ-glycerophosphate concentrations and mineralization by both osteoblastsand fibroblasts. Increased mineralization by ligament fibroblasts isobserved with increasing concentration of β-glycerophosphate, a mediaadditive commonly used in osteoblast cultures. These reports stronglysuggest the plasticity of the fibroblast response and that thede-differentiation of ligament fibroblasts is a function of mineralcontent in vitro.

Progressing through the four different zones which make up the nativeACL insertion zone, several cell types are identified: ligamentfibroblasts, chondrocytes, hypertrophic chondrocytes and osteoblasts,osteoclasts, and osteocytes. The development of in vitro multi-cell typeculture systems facilitates the formation of the transition zones.

No reported studies on either the co-culture of ligament fibroblastswith osteoblasts, nor on the in vitro and in vivo regeneration of thebone-ligament interface are known.

No reported studies which examine the potential of multi-phasedscaffolds in facilitating the fixation of ligament or tendon to bone areknown. As the interface between graft and bone is the weakest pointduring the initial healing period, recent research efforts in ACL tissueengineering have concentrated on design of multi-phased scaffolds inorder to promote graft integration.

Goulet et al. (2000) developed a bio-engineered ligament model, whereACL fibroblasts were added to the structure and bone plugs were used toanchor the bioengineered tissue. Fibroblasts isolated from human ACLwere grown on bovine type I collagen, and the bony plugs were used topromote the anchoring of the implant within the bone tunnels.

Cooper et al. (2000) and Lu et al. (2001) developed a tissue engineeredACL scaffold using biodegradable polymer fibers braided into a 3-Dscaffold. This scaffold has been shown to promote the attachment andgrowth of rabbit ACL cells in vitro and in vivo. However, no multiphasedscaffolds for human ligament-to-bone interface are known.

SUMMARY

This application describes scaffold apparatuses for musculoskeletaltissue engineering.

A scaffold apparatus, according to one preferred embodiment, ismulti-phasic and can support growth, maintenance and differentiation ofmultiple tissue and cell types. The multi-phasic scaffold apparatus hasa gradient of calcium phosphate content across the phases, and isbiomimetic, biodegradable and/or osteointegrative.

A scaffold apparatus, according to another embodiment, includesmicrospheres of selected sizes and/or composition. The microspheres arelayered to have a gradient of microsphere sizes and/or compositions. Thescaffold provides a functional interface between multiple tissue types.

A multi-phased scaffold apparatus for providing a functional interfacebetween bone and soft tissue is also described. The multi-phasedscaffold apparatus, according to one embodiment, includes microspheresas one phase of the scaffold, and a mesh as another phase of thescaffold. The microspheres and the mesh are sintered together.

A scaffold apparatus for soft tissue-to-bone interface tissueengineering is also described. The apparatus comprises four (or more)regions. The first region comprises composite microspheres of a firstsize and composition optimized to promote growth, proliferation, anddifferentiation of a first cell type for integration and growth of afirst tissue type. The second region is joined to the first region, andcomprises microspheres and/or a fibrous mesh which have a second sizeand a second composition. The third region is joined to the secondregion, and comprises microspheres and/or a fibrous mesh which has athird size and a third composition. The second and third regions areoptimized to promote growth, proliferation and differentiation of asecond cell type for integration and formation of a second tissue type.The fourth region is joined to the third region, and comprisesmicrospheres and/or a fibrous mesh which have a composition adapted topromote growth, proliferation, and differentiation of a third cell typefor integration and growth of a third tissue type.

This application also describes methods for preparing a scaffold formusculoskeletal tissue engineering. According to one exemplaryembodiment, a method for preparing a scaffold comprises (a) processing aplurality of microspheres, including incorporating calcium phosphate(and/or bioglass) into the microspheres, (b) laying the processedmicrospheres in a mold, the microspheres in the mold presenting agradient of microsphere sizes and/or compositions, and (c) sinteringtogether the microspheres in the mold above the polymer glass transitiontemperature.

According to another embodiment, a method for preparing a multi-phasescaffold for musculoskeletal tissue engineering. The method, accordingto one embodiment, comprises (a) processing a plurality of microspheres,including incorporatingcalcium phosphate (and/or bioglass) into themicrospheres, (b) laying the processed microspheres in a mold, whereinthe microspheres in the mold present a gradient of microsphere sizes fora first phase and a second phase of the multi-phase scaffold, withmicrospheres of the first phase being in a first range of sizes, andwith microspheres of the second phase being in a second range of sizeslarger than the first range of sizes, (c) sintering together themicrospheres in the mold above a glass transition temperature, and (d)sintering a fiber mesh, as a third phase of the multi-phase scaffold,onto the microsphere construct prepared in (c).

According to another exemplary embodiment, a method for preparing amulti-phase scaffold for musculoskeletal tissue engineering, cancomprise the steps of (a) forming a mesh scaffold by sintering togethera plurality of mesh segments as a first phase of the multi-phasescaffold, (b) forming a second scaffold by sintering together aplurality of poly-lactide-co-glycolide microspheres as a second phase ofthe multi-phase scaffold, (c) forming a third scaffold by sinteringtogether a plurality of microspheres formed of a composite ofpoly-lactide-co-glycolide and bioactive glass as a third phase of themulti-phase scaffold, and (d) sintering together said mesh scaffold,said second scaffold and said third scaffold.

This application also describes methods for producing polymer/ceramiccomposite microspheres.

BRIEF DESCRIPTION OF THE FIGURES

FIG. 1A shows a schematic diagram of a scaffold apparatus, according toone embodiment.

FIG. 1B shows a schematic diagram of a scaffold apparatus, according toanother embodiment

FIG. 2 shows a schematic diagram of a multi-phased scaffold apparatus,according to another embodiment.

FIG. 3 shows a flow chart for a method for preparing a scaffold,according to an exemplary embodiment.

FIG. 4 shows a flow chart for a method for preparing a multi-phasedscaffold, according to another embodiment.

FIG. 5A: Posterior view of an intact bovine anterior cruciate ligament(ACL) connecting the femur to the tibia (left). FIG. 5B: Environmentalscanning electron microscope (ESEM) image of transition from ligament(L) to fibrocartilage (FC) to bone (B) at the ACL insertion (upperright). FIG. 5C: Histological micrograph of similar ACL to boneinterface additionally showing mineralized fibrocartilage (MFC) zone(lower right).

FIGS. 6A and 6B show Bovine tibial-femoral joint after ACL and insertionsite extraction (right), ACL and insertion sites after excision.

FIG. 7A shows FTIR Spectra of BG immersed in SBF for up to 7 days.Presence of an amorphous Ca-P layer at 1 day, and of a crystalline layerat 3 days.

FIG. 7B: SEM image of Ca-P nodules on BG surface (3 days in SBF).Nodules are ˜1 μm in size initially, and grew as immersion continued(15,000×). FIG. 7C: EDXA spectrum of BG surfaces immersed in SBF for 3days. The relative Ca/P ratio is ≈1.67.

FIGS. 8A and 8B show environmental SEM images of Bovine ACL insertionSite (1 and 2), including a cross section of the ACL-femur insertionsite, ACL fiber (L) left, fibrocartilage region (FC) middle, andsectioned bone (B) right (FIG. 8A: 250×; FIG. 8B: 500×).

FIG. 9A: SEM of the cross section of the femoral insertion zone, 1000×;FIG. 9B: EDAX of the femoral insertion zone. The peak intensities of Ca,P are higher compared to those in ligament region.

FIG. 10 shows apparent modulus versus indentation X-position acrosssample.

FIGS. 11A and 11B show X-Ray CT scans of discs made ofpoly-lactide-co-glycolide (PLAGA) 50:50 and bioactive glass (BG)submerged in SBF for 0 days (FIG. 11A) and days; FIG. 11B shows theformation of Ca-P over time.

FIG. 12A: SEM image; FIG. 12B: EDAX of PLAGA-BG immersed in SBF for 14days.

FIG. 13 shows osteoblast grown on PLAGA-BG, 3 weeks.

FIG. 14 shows higher type I collagen type synthesis on PLAGA-BG.

FIG. 15A: ALZ stain, ACL fibroblasts 14 days, 20×; FIG. 15B: ALZ stain,interface, ACL 14 days, 20×; FIG. 15C: ALZ stain, osteoblasts, ACL 14days, 20×.

FIG. 16A: ALP stain, ACL fibroblasts, 7 days, 32×; FIG. 16B: ALP+DAPIstain, co-culture, 7 days, 32×; FIG. 16C: ALP stain, osteoblasts, 7days, 32×.

FIGS. 17A-17F show images of multiphase scaffold (FIGS. 17A-17C) andblow-ups of respective sections (FIGS. 17D-17F).

FIGS. 18A-18C show multiphasic scaffold for co-culture of ligamentfibroblasts and osteoblasts; FIG. 18A and FIG. 18B: images of a samplescaffold; FIG. 18C: schematic of scaffold design depicting the threelayers.

FIGS. 19A-19D show Micromass co-culture samples after 14 days. FIG. 19A:H&E stain; FIG. 19B: Alcian blue; FIG. 19C: Type I collagen (green);FIG. 19D: Type II collagen (green)+Nucleic stain (red).

FIGS. 20A and 20B show RT-PCR gel for day 7 micromass samples. FIG. 20A:Type X collagen expression. FIG. 20B: Type II collagen expression. (C:control micromass sample; E: experimental co-culture sample).

FIGS. 21A and 21B show SEM image of cellular attachment to PLAGA-BGscaffold after 30 min; FIG. 21A: chondrocyte control (2000×); FIG. 21B:co-culture (1500×).

FIGS. 22A-22C show Cellular attachment to PLAGA-BG scaffold; FIG. 22A:chondrocyte control, day 1 (500×); FIG. 22B: co-culture, day 1 (500×).FIG. 22C: co-culture, day 7 (750×).

FIG. 23 shows multi-phased scaffold, according to one embodiment (for invitro co-culture).

FIG. 24 shows a schematic diagram depicting a fabrication process of acomposite (PLAGA-BG) of PLAGA and BG, in thin film form and as a 3-D,porous scaffold.

FIG. 25 shows a tabular summary of results from structuralcharacterizations of an as-fabricated composite scaffold.

FIG. 26 shows EDXA spectra of the PLAGA-BG composite immersed in a SBFfor 14 days.

FIG. 27 shows a comparison of the expression of type I collagen by humanosteoblast-like cells cultured on the PLAGA-BG composite versus on TCPScontrols and on PLAGA alone.

FIG. 28-1 shows a table of porosimetry data, including intrusion volume,porosity, and pore diameter data, in another set of experiments.

FIGS. 28-2A through 28-2C show fluorescence microscopy images (day 28,x10) for Phases A through C, respectively.

FIGS. 28-3A and 28-3B are images showing extracellular matrix productionfor Phases B and C, respectively.

FIGS. 29-1A through 29-1D show SEM images, in another set ofexperiments:

A) Phase C, Day 0—×1000;

B) Phase C, Day 28—×1000;

C) Phase A, Day 28—×1000; and

D) Phase B, Day 28—×70.

FIGS. 29-2A through 29-2F show fluorescence microscopy images:

A) Phase A, Day 0, ×10;

B) Phase B, Day 0, ×10;

C) Phase C, Day 0, ×10;

D) Phase A, Day 28, ×10;

E) Phase B, Day 28, ×10; and

F) Phase C, Day 28, ×10.

FIGS. 29-3A1 through 29-3A3 show Trichrome images (Day 0, ×10) of PhaseA, Phase B and Phase C, respectively.

FIGS. 29-3B1 and 29-3B2 show Picrosirius Red images of Phase B and PhaseC, respectively.

FIGS. 29-3C shows a von Kossa image of Phase C.

FIGS. 30-1 a through 30-1 f show images of osteoblast and fibroblast inculture, in another set of experiments:

a) Day 0, 5×;

b) Day 0, 5×;

c) Day 1, 5×;

d) Day 2, 5×;

e) Day 1, 32× (cell contact); and

f) Day 1, 32×.

FIGS. 30-2 a through 30-2 c show stained images:

a) live-dead stain of 1 hr sample, 5×;

b) ALP stain of ob and fb, day 2, 20×; and

c) collagen I staining, day 6, 20×.

FIG. 31-1 shows a schematic of the experimental design, in another setof experiments, for in vitro evaluations of human osteoblasts andfibroblasts co-cultured on multi-phased scaffolds.

FIG. 31-2 shows a graph which demonstrates cell proliferation in PhasesA, B, and C during 35 days of human hamstring tendon fibroblast andosteoblast co-culture on multiphased scaffolds.

FIGS. 31-3A and 31-3B graphically show Mechanical testing data formultiphased scaffolds seeded with human hamstring tendon fibroblasts andhuman osteoblasts over 35 days of culture (n=4).

FIG. 32-1 schematically shows a method for producing multi-phasicscaffolds, in another set of experiments. First, Ethicon PLAGA mesh iscut into small pieces and inserted into a mold. By applying compressionforce (F) and heating (H) at 150° C. for time (t)=20 hours, the meshsegments are sintered into a mesh scaffold, which is removed from themold. Next, PLAGA microspheres are inserted into the mold, sintered,then removed as a second scaffold. The same process is performed for thePLAGA-BG microspheres. Finally, Phases A and B are joined by solventevaporation, then all three scaffolds are inserted into the mold andsintered together, forming the final multi-phasic scaffold.

FIG. 32-2 shows a schematic of a co-culture experimental design.

FIG. 32-3 shows a table summarizing mercury porosimetry data.

FIGS. 32-4A and 32-4B show graphically scaffold phase thicknesses anddiameters, in the experiments of FIG. 32-1 through FIG. 32-3.

FIG. 32-5 shows graphically a comparison of microsphere initial mass andfinal mass after undergoing a sintering process.

FIGS. 32-6A and 32-6B show graphically mechanical testing data formultiphased scaffolds seeded with human hamstring tendon fibroblasts andhuman osteoblasts over days of culture (n=4). Scaffolds were tested inuniaxial compression. Compressive modulus (A) and yield strength (B)were calculated from the resulting stress-strain curves. Both cellseeded (C) and acellular (AC) scaffolds were examined at days 0, 7, 21,and 35. Scaffold compressive modulus was significantly greater at day 0than for all subsequent time points and groups (p<0.05).

FIG. 33-1 shows a table illustrating the compositions of polymersolutions tested, in another set of experiments.

FIG. 33-2 shows a table illustrating drum rotational velocity (rpm) andsurface velocity (m/s) for each gear.

FIGS. 33-3A and 33-3D show SEMS of electrospun meshes spun at:

A) 1^(st) gear, 7.4 m/s;

B) 2^(nd) gear, 9.4 m/s;

C) 3^(rd) gear, 15 m/s; and

D) 4^(th) gear, 20 m/s.

FIGS. 33-4A and 33-4B show scanning electron microscopy (SEM) images ofanother embodiment of multi-phased scaffold, with 85:15 PLAGAelectrospun mesh joined with PLAGA:BG composite microspheres.

FIG. 34 schematically shows one exemplary embodiment of multi-phasedscaffold as a hamstring tendon graft collar which can be implementedduring ACL reconstruction surgery to assist with hamstringtendon-to-bone healing.

FIG. 35 shows a flow chart for a method for preparing a scaffold,according to another exemplary embodiment.

DETAILED DESCRIPTION

In order to facilitate an understanding of the material which follows,one may refer to Freshney, R. Ian. Culture of Animal Cells—A Manual ofBasic Technique (New York: Wiley-Liss, 2000) for certain frequentlyoccurring methodologies and/or terms which are described therein.

However, except as otherwise expressly provided herein, each of thefollowing terms, as used in this application, shall have the meaning setforth below.

As used herein, “bioactive” shall include a quality of a material suchthat the material has an osteointegrative potential, or in other wordsthe ability to bond with bone. Generally, materials that are bioactivedevelop an adherent interface with tissues that resist substantialmechanical forces.

As used herein, “biomimetic” shall mean a resemblance of a synthesizedmaterial to a substance that occurs naturally in a human body and whichis not rejected by (e.g., does not cause an adverse reaction in) thehuman body.

As used herein, “chondrocyte” shall mean a differentiated cellresponsible for secretion of extracellular matrix of cartilage.

As used herein, “fibroblast” shall mean a cell of connective tissue,mesodermally derived, that secretes proteins and molecular collagenincluding fibrillar procollagen, fibronectin and collagenase, from whichan extracellular fibrillar matrix of connective tissue may be formed.

Generally, “glass transition temperature” is the temperature at which,upon cooling, a noncrystalline ceramic or polymer transforms from asupercooled liquid into a rigid glass. The noncrystalline ceramic orpolymer may be of multiple form and composition, and may be formed asmicrospheres. In the context of a sintering process, such as discussedin this application, the polymer chains from adjacent microspherestypically entangle, effectively forming a bond between the microspheresupon cooling. As the polymer is heated above its glass transitiontemperature, long range polymer chain motion begins.

As used herein, “graft fixation device” shall mean a device for fixationof a graft, including but not limited to staples, interference screwswith or without washers, press fit EndoButton® devices and Mitek® Anchordevices.

As used herein, “interference screw” shall mean a device indicated forsoft tissue-bone fixation. The device may be used in, for example,anterior cruciate ligament surgery. The device may include, but is notlimited to, at least titanium cannulated interference screws, PLLAabsorbable interference screws, and Poly-L-Lactide interference screws.

As used herein, “matrix” shall mean a three-dimensional structurefabricated from biomaterials. The biomaterials can bebiologically-derived or synthetic.

As used herein, “osteoblast” shall mean a bone-forming cell that isderived from mesenchymal osteoprognitor cells and forms an osseousmatrix in which it becomes enclosed as an osteocyte. The term is alsoused broadly to encompass osteoblast-like, and related, cells, such asosteocytes and osteoclasts.

As used herein, “osteointegrative” shall mean ability to chemically bondto bone.

As used herein, “polymer” shall mean a chemical compound or mixture ofcompounds formed by polymerization and including repeating structuralunits. Polymers may be constructed in multiple forms and compositions orcombinations of compositions.

As used herein, “porosity” shall mean the ratio of the volume ofinterstices of a material to a volume of a mass of the material.

As used herein, “particle reinforcement” shall mean a process forforming a composite with a higher strength than the original material(for example, a polymer) by adding particles of a reinforcing materialwith a higher strength (for example, a ceramic).

As used herein, “sintering” shall mean densification of a particulatepolymer compact involving a removal of pores between particles (whichmay be accompanied by equivalent shrinkage) combined with coalescenceand strong bonding between adjacent particles. The particles may includeparticles of varying size and composition, or a combination of sizes andcompositions.

This application describes scaffolds having a gradient of properties(such as structural properties, pore diameter, chemical properties,mechanical properties, etc.), for the repair of musculoskeletal tissue.The scaffold is preferably multi-phased, biodegradable, andosteointegrative.

The following exemplary embodiments and experimental details sectionsare set forth to aid in an understanding of the subject matter of thisdisclosure but are not intended to, and should not be construed to,limit in any way the invention as set forth in the claims which followthereafter.

A scaffold apparatus, according to one preferred embodiment, ismulti-phasic, including first, second and third phases, and preferablycan support growth, maintenance and differentiation of multiple tissueand cell types.

The first phase comprises a first material adapted for integration andgrowth (for example, by including one or more osteogenic agents,osteogenic materials, osteoinductive agents, osteoinductive materials,osteoconductive agents, osteoconductive materials, growth factors,chemical factors, etc.) of a first tissue type and is seeded with afirst type of cells (for example, osteoblasts, osteoblast-like cells,stem cells, etc.). The material of the first phase may include, but isnot limited to, microspheres, foams, sponges and any other threedimensional (3-D) scaffold construct consisting of polymer and/orceramic. Polymers may include, but is not restricted to, anybiodegradable polymer such as any of the poly-(α-hydroxy acids), ornatural polymers such as silk, collagen, or chitosan. Ceramics mayinclude but are not limited to bioactive glass, hydroxyapatite, betatricalcium phosphate, or any other calcium phosphate material.

The third phase comprises a second material adapted for integration andgrowth of a second tissue type seeded with a second type of cells (forexample, fibroblasts, chondrocytes, stem cells, etc.). The third phasemay include a composite of materials, including, but not limited to,microspheres, a fiber mesh, degradable polymers, etc.

The second phase is an interfacial zone between the first and thirdphases.

The multi-phasic scaffold apparatus preferably has a gradient of calciumphosphate content across the phases, and is preferably biomimetic,biodegradable (that is, each phase is degradable) and/orosteointegrative.

A scaffold apparatus for musculoskeletal tissue engineering, accordingto another embodiment, may include microspheres of selected sizes and/orcomposition. The microspheres may be layered to have a gradient ofmicrosphere sizes and/or compositions. The scaffold may provide afunctional interface between multiple tissue types (for example, softtissue and bone).

FIG. 1A shows schematically a multi-phased scaffold apparatus 10comprising phase A, phase B, and phase C. Phases A-C have a gradient ofproperties. The gradient of properties across phases A-C of the scaffoldmay include mineral content (for example, Ca-P), mechanical properties,chemical properties, structural properties, porosity, geometry, etc. Itshould be apparent to one skilled in the art that although apparatus 10has three phases, the apparatus can be integrated in a scaffold withfour or more phases.

For example, the multi-phased scaffold may contain a gradient of Ca-Pconcentrations. Phase A may be constructed of fiber mesh with alignedfibers and with no Ca-P, phase C may be constructed of polymer-ceramiccomposite with high Ca-P, and phase B may be constructed ofpolymer-ceramic composite with lower Ca-P than phase C.

The scaffold apparatus can promote growth and maintenance of multipletissue types. The scaffold may support growth, maintenance anddifferentiation of multiple tissue and cell types. The multi-phasedscaffold may mimic the inhomogeneous properties of the insertion zonebetween soft tissue and bone, resulting in desired growth, phenotypicexpression, and interactions between relevant cell types.

The phases of the scaffold may be inhomogeneous in properties. Thephases may have zonal differences in mineral content and matrixmorphology designed to mimic the tissue-bone interface and to facilitatethe growth and maintenance of different tissues. The phases may differin morphology. For example, phase A can include a porous fibrous mesh,while phases B and C include microspheres. According to anotherembodiment, the scaffold may include a composite of microspheres and afiber mesh.

The scaffold preferably includes multiple phases. According to oneembodiment, one phase (for example, phase A) supports growth andmaintenance of soft tissue, another phase (for example, phase C)supports growth and maintenance of bone, and a third phase is aninterfacial zone between the first and second phases. The first phasefor supporting growth and maintenance of the soft tissue may be seededwith at least one of fibroblasts, chondrocytes and stem cells. Thesecond phase for supporting growth and maintenance of the bone may beseeded with at least one of osteoblasts, osteoblast-like cells and stemcells. The second phase can contain at least one of osteogenic agents,osteogenic materials, osteoinductive agents, osteoinductive materials,osteoconductive agents, osteoconductive materials, growth factors andchemical factors.

Further, at least one of said first phase and said second phase may beseeded with one or more agents by using a microfluidic system.

The third phase may include some of the microspheres. The third phasecan include a gradient of microsphere sizes and/or a gradient ofmicrosphere compositions. The microspheres in the third phase may bejoined by sintering in at least one stage.

The second phase may include additional microspheres. The second phasecan comprise one of polymeric and composite microspheres including arange of diameters or a gradient of diameter. At least some of themicrospheres of the third phase may be in a first range of sizes, andthe additional microspheres of the second phase may be in a second rangeof sizes lower than the first range of sizes.

The second phase can comprise polymeric hydrogels of one of polyethyleneglycol and hydroxyethyl methacrylate. The hydrogel may comprise one ormore of poly(ethylene glycol), agarose, alginate, 2-hydroxyethylmethacrylate and polyacrylamide. The second phase can comprise collagengels with varied mineral content.

The scaffold may include a composite of microspheres and a fiber mesh.The fiber mesh may be a degradable polymer. For example, the first phasemay include a fiber mesh. The fiber mesh of the first phase and themicrospheres of the third phase may be sintered together. The fiber meshmay be electrospun.

The mesh can include one or more desired agents and/or compound. Forexample, at least one of bioactive agents and peptides may coat thesurface of the mesh. The bioactive agents and peptides can enhancedifferentiation, proliferation and attachment of cells and specific celltypes. Also or alternatively, at least one of bioactive agents andpeptides can directly be incorporated into the mesh.

According to one embodiment, the scaffold may include multiple phasesjoined by a gradient of properties. The multiple phases joined by thegradient of properties may be processed through one or more sinteringstages. The gradient of properties across the multiple phases of thescaffold can include mechanical properties, chemical properties, mineralcontent, structural properties, porosity and/or geometry.

The scaffold apparatus can include plural phases of microspheres. Forexample, a first phase of the microspheres can comprise polymer and asecond phase of the microspheres can comprise one of bioactive glass andcalcium phosphate. Varying concentrations of calcium phosphate can beincorporated into the microspheres. The calcium phosphate can beselected from a group comprising tricalcium phosphate, hydroxyapatite,and a combination thereof. The polymer can be selected from a groupcomprising aliphatic polyesters, poly(amino acids),copoly(ether-esters), polyalkylenes oxalates, polyamides,poly(iminocarbonates), polyorthoesters, polyoxaesters, polyamidoesters,poly(ε-caprolactone)s, polyanhydrides, polyarylates, polyphosphazenes,polyhydroxyalkanoates, polysaccharides, and biopolymers, and a blend oftwo or more of the preceding polymers. The polymer can comprise at leastone of poly(lactide-co-glycolide), poly(lactide) and poly(glycolide).

The microspheres may comprise one or more of CaP, bioactive glass,polymer, etc. The microspheres may be processed through one or moresintering stages.

The microspheres may comprise one or more desired agents or compounds.For example, at least one of bioactive agents and peptides may coat thesurface of at least some of the microspheres. The bioactive agents andpeptides can enhance at least one of differentiation, proliferation andattachment of cells and specific cell types. Also or alternatively, atleast one of bioactive agents and peptides can directly be incorporatedinto at least some of the microspheres. The microspheres canadditionally include one or more agents selected from a group comprisingantiinfectives, hormones, analgesics, anti-inflammatory agents, growthfactors, chemotherapeutic agents, anti-rejection agents and RGDpeptides.

The apparatus is preferably biomimetic, biodegradable and/orosteointegrative.

According to one exemplary embodiment, the apparatus may be integratedin a graft fixation device. The graft fixation device may be used, forexample, for graft fixation at the bone tunnels during anterior cruciateligament reconstruction.

According to another embodiment, the apparatus may be integrated in aninterference screw.

In addition, the scaffold apparatus, according to another exemplaryembodiment, may be integrated in a graft collar. The graft collar hasmany applications. For example, the graft collar may be adapted forhamstring tendon-to-bone healing. As another example, the graft collarcan be adapted for peridontal ligament repair. Further, the graft collarmay be adapted for spinal repair.

A scaffold apparatus for soft tissue-to-bone interface tissueengineering, according to another exemplary embodiment, is shownschematically in FIG. 1B. Apparatus 15 includes a first region H, asecond region I which is joined to region H, a third region J which isjoined to region I, and a fourth region K which is joined to region J.

Region H comprises composite microspheres of a first size andcomposition optimized to promote growth, proliferation, anddifferentiation of a first cell type for integration and growth of afirst tissue type. The composite microspheres of region H can include arange of sizes.

Region I comprises at least one of microspheres and a fibrous meshhaving a second size and a second composition. The microspheres and/orfibrous mesh of region I can include a range or gradient of sizes,and/or a gradient of compositions.

Region J comprises at least one of a microsphere and a fibrous meshhaving a third size and a third composition. Regions I and J areoptimized to promote growth, proliferation and differentiation of asecond cell type for integration and formation of a second tissue type.The microspheres and/or fibrous mesh of region J can include a range orgradient of sizes, and/or a gradient of compositions.

Region K comprises at least one of a microsphere and a fibrous meshhaving a composition adapted to promote growth, proliferation, anddifferentiation of a third cell type for integration and growth of athird tissue type. The fibrous mesh may be electrospun.

The regions H-K can be joined together through one of a solid statesintering process and a solvent aggregation process, in which selectedgrowth factors or bioactive agents are incorporated into each region topromote formation, growth and integration of said first, second andthird types of tissues. The scaffold apparatus may be integrated in agraft collar.

A multi-phased scaffold apparatus for providing a functional interfacebetween bone and soft tissue, according to an embodiment schematicallyshown in FIG. 2, includes microspheres as one phase, and a mesh asanother phase. The microspheres and the mesh may be sintered together.

FIG. 2 shows schematically a multi-phased scaffold apparatus 20comprising phase X and phase Y. Microspheres may be one phase of thescaffold and a mesh may be another phase of the scaffold. Themicrospheres and the mesh may be sintered together. The apparatus 20 maybe integrated in a scaffold which includes multiple phases (for example,three or more).

The microsphere and mesh structure of the scaffold may be geometricallyheterogeneous, including a fiber mesh for culturing fibroblasts and anopen-pore structure for osteoblasts. At least one zone of hydrogels oropen-pore structure for chondrocytes may also be included. Themicrosphere and mesh components may be incorporated into the scaffold toallow for the co-culturing of multiple cell types to mimic the multitudeof cell types found at native tissue interfaces. The mesh can beelectrospun.

The scaffold may be modified to achieve specific cell cultureparameters, for example, by including microspheres of varying diametersto vary the porosity of the scaffold in different regions. Furthermore,the scaffold may be fabricated in a variety of geometries. For example,the scaffold apparatus can be integrated in a graft collar.

This application also describes methods for preparing a scaffold formusculoskeletal tissue engineering. A method, according to oneembodiment (FIG. 3), includes (a) processing a plurality of microspheres(step S31), including incorporating calcium phosphate into themicrospheres, (b) laying the processed microspheres in a mold (stepS33), the microspheres in the mold presenting a gradient of microspheresizes and/or compositions, and (c) sintering together the microspheresin the mold above a glass transition temperature (step S35).

Additional steps may optionally be added to the method to impartadditional scaffold features or characteristics. For example, the methodmay further include sintering a fiber mesh onto the microsphereconstruct to provide a functional interface between multiple tissuetypes. Further, the method may further comprise electrospinning saidfiber mesh prior to attaching the electrospun fiber mesh onto themicrosphere construct.

Varying concentrations of calcium phosphate may be incorporated into themicrospheres. The calcium phosphate incorporated into the microspheresmay include hydroxyapatite, tricalcium phosphate, etc.

The particulate phase of the microspheres may include bioactive glass.Varying porosity or concentrations of bioactive glass may beincorporated into the microspheres.

The method may further include applying a particle reinforcement processto the microspheres. The method may further include incorporatingparticulates in the microspheres prior to the sintering step tostrengthen the microspheres.

A method for preparing a multi-phase scaffold for musculoskeletal tissueengineering, according to an exemplary embodiment (FIG. 4), includes (a)processing a plurality of microspheres (step S41), includingincorporating calcium phosphate into the microspheres, (b) laying theprocessed microspheres in a mold (step S43), wherein the microspheres inthe mold presenting a gradient of microsphere sizes for a first phaseand a second phase of the multi-phase scaffold, with microspheres of thefirst phase being in a first range of sizes, and with microspheres ofthe second phase being in a second range of sizes larger than the firstrange of sizes, (c) sintering together the microspheres in the moldabove a glass transition temperature (step S45), and (d) sintering afiber mesh, as a third phase of the multi-phase scaffold, onto themicrosphere construct prepared in (c) (step S47).

Additional steps may optionally be included. For example, the method mayfurther include seeding the third phase with at least one of fibroblasts(for example, human hamstring tendon fibroblasts), chondrocytes and stemcells. The seeding of the third phase supports growth and maintenance ofsoft tissue. Also, the method can include seeding the first phase withat least one of osteoblasts and stem cells. The seeding of the firstphase supports growth and maintenance of bone. The method may furtherinclude seeding the second phase with at least one of chondrocytes andstem cells. Seeding of the second phase can support growth andmaintenance of fibrocartilage.

The first phase may support growth and maintenance of bone. The thirdphase may support growth and maintenance of soft tissue. The secondphase may serve at least as an interfacial zone between the first phaseand the third phase.

For example, the method may further comprise seeding the first phasewith first cells, for supporting growth and maintenance of the bone,seeding the third phase with second cells for supporting growth andmaintenance of the soft tissue, and allowing at least some of said firstcells and said second cells to migrate to the second phase.

In addition, the method may further comprise seeding at least one ofsaid first, second and phases with one or more agents by using amicrofluidic system.

Further, the method may further comprise electrospinning said fiber meshprior to attaching the fiber mesh onto the microsphere construct.

This application also provides methods for producing polymer/ceramiccomposite microspheres. The composite microspheres can be formed byapplying an emulsion and solvent evaporation process. The compositemicrospheres can comprise a degradable polymer and one of bioactiveglass and calcium phosphate ceramics. The degradable polymer can bedissolved in a solvent. The bioactive glass and/or calcium phosphateceramics can be mixed into the polymer solution. A suspension of thebioactive glass and/or calcium phosphate ceramics in the polymersolution can be poured into a stirring surfactant solution.

The degradable polymer may be a polymer selected from the groupconsisting at least of aliphatic polyesters, poly(amino acids),copoly(ether-esters), polyalkylenes oxalates, polyamides,poly(iminocarbonates), polyorthoesters, polyoxaesters, polyamidoesters,poly(e-caprolactone)s, polyanhydrides, polyarylates, polyphosphazenes,polyhydroxyalkanoates, polysaccharides and biopolymers.

Calcium phosphate and/or bioactive glass particles may be encapsulatedin the microspheres during emulsion.

A method, according to another exemplary embodiment (FIG. 35), forpreparing a multi-phase scaffold for musculoskeletal tissue engineering,can comprise the steps of (a) forming a mesh scaffold by sinteringtogether a plurality of mesh segments as a first phase of themulti-phase scaffold (step S351), (b) forming a second scaffold bysintering together a plurality of poly-lactide-co-glycolide microspheresas a second phase of the multi-phase scaffold (step S352), (c) forming athird scaffold by sintering together a plurality of microspheres formedof a composite of poly-lactide-co-glycolide and bioactive glass as athird phase of the multi-phase scaffold (step S353), and (d) sinteringtogether said mesh scaffold, said second scaffold and said thirdscaffold (step S354). Steps S351 through S353 may be performed in anyorder.

The specific embodiments described herein are illustrative, and manyvariations can be introduced on these embodiments without departing fromthe spirit of the disclosure or from the scope of the appended claims.Elements and/or features of different illustrative embodiments may becombined with each other and/or substituted for each other within thescope of this disclosure and appended claims.

Further non-limiting details are described in the following ExperimentalDetails section which is set forth to aid in an understanding of theinvention but is not intended to, and should not be construed to, limitin any way the claims which follow thereafter.

Experimental Details First Set of Experiments

To address the challenge of graft fixation to subchondral bone, a normaland functional interface may be engineered between the ligament andbone. This interface, according to one exemplary embodiment, wasdeveloped from the co-culture of osteoblasts and ligament fibroblasts ona multi-phased scaffold system with a gradient of structural andfunctional properties mimicking those of the native insertion zones toresult in the formation of a fibrocartilage-like interfacial zone on thescaffold. Variations in mineral content from the ligament proper to thesubchondral bone were examined to identify design parameters significantin the development of the multi-phased scaffold. Mineral content (Ca-Pdistribution, Ca/P ratio) across the tissue-bone interface wascharacterized. A multi-phased scaffold with a biomimetic compositionalvariation of Ca-P was developed and effects of osteoblast-ligamentfibroblast co-culture on the development of interfacial zone specificmarkers (proteoglycan, types II & X collagen) on the scaffold wereexamined.

The insertion sites of bovine ACL to bone (see FIGS. 5A-5C) wereexamined by SEM. Pre-skinned bovine tibial-femoral joints were obtained.The intact ACL and attached insertion sites were excised with a scalpeland transferred to 60 mm tissue culturing dishes filled with Dulbecco'sModified Eagle Medium (DMEM) (see FIGS. 6A and 6B). After isolation, thesamples were fixed in neutral formalin overnight, and imaged byenvironmental SEM (FEI Quanta Environmental SEM) at an incident energyof 15 keV. ACL attachment to the femur exhibited an abrupt insertion ofthe collagen bundle into the cartilage/subchondral bone matrix.Examination of collagen bundle revealed that the surface was ruffled andsmall collagen fibrils can be seen. When a cross section was imaged,three distinct zones at the insertion site were evident: ligament (L),fibrocartilage (FC), and subchondral bone (B). The interface regionspans proximally 200 μm. These cross section views showed the insertionof Sharpey fiber into the fibrocartilage (see FIGS. 7A, 7B and 7C).Mineralized fibrocartilage was not distinguishable with regularcartilage from these images.

The insertion sites of bovine ACL to bone were examined by scanningelectron microscopy (SEM). Bovine tibial-femoral joints were obtained.The intact ACL and attached insertion sites were excised with a scalpeland transferred to 60 mm tissue culturing dishes filled with Dulbecco'sModified Eagle Medium (DMEM). After isolation, the samples were fixed inneutral formalin overnight, and imaged by environmental SEM (FEI QuantaEnvironmental SEM) at 15 keV.

ACL attachment to the femur exhibited an abrupt insertion of thecollagen bundle into subchondral bone. When a cross section was imaged(see FIGS. 8A and 8B), three distinct zones at the insertion site wereevident: ligament (L), fibrocartilage (FC), and subchondral bone (B).Sharpey fiber insertion into the fibrocartilage (see FIG. 8A) wasobserved. The bovine interface region spans proximally 600 μm.Examination of the interface using energy dispersive X-ray analysis(EDAX, FEI Company) enable the mineralized and non-mineralized FC zonesto be distinguished. A zonal difference in Ca and P content was measuredbetween the ligament proper and the ACL-femoral insertion (see Table I).

TABLE I Region Ca/P Analyzed Ca P Ratio S Ligament 1.69 2.98 0.57 3.71Insertion 5.13 5.93 0.87 19.50

At the insertion zone (see FIGS. 9A and 9B), higher Ca and P peakintensities were observed, accompanied by an increase in Ca/P ratio ascompared to the ligament region. Higher sulfur content due to thepresence of sulfated proteoglycans at the FC region was also detected.The zonal difference in Ca-P content was correlated with changes instiffness across the interface. Nanoindentation measurements wereperformed using atomic force microscopy (AFM, Digital Instruments). Anincreasing apparent modulus was measured as the indentation testingposition moved from the ligament region into the transition zone (seeFIG. 10).

Ca-P distribution on polylactide-co-glycolide (50:50) and 45S5 bioactiveglass composite disc (PLAGA-BG) after incubation in a simulated bodyfluid (SBF) was evaluated using μCT (μCT 20, Scanco Medical,Bassersdorf, Switzerland) following the methods of Lin et al. The samplewas loaded into the system, scanned at 20 mm voxel resolution and anintegration time of 120 ms. FIGS. 11A and 11B compare the amount ofcalcified region (dark areas) observed on the PLAGA-BG disc as afunction of incubation time in SBF (from day 0 to day 28). Using customimage analysis software, it was determined that at day 0, themineralized region corresponded to 0.768% of the total disc (quartered)area, and at day 28, the mineralized region corresponded to 12.9% of thetotal area. Results demonstrate the Ca-P distribution on scaffoldsmeasured by μCT analysis.

The scaffold system developed for the experiments was based on a 3-Dcomposite scaffold of ceramic and biodegradable polymers. A compositesystem has been developed by combining poly-lactide-co-glycolide (PLAGA)50:50 and bioactive glass (BG) to engineer a degradable,three-dimensional composite (PLAGA-BG) scaffold with improved mechanicalproperties. This composite was selected as the bony phase of themulti-phased scaffold as it has unique properties suitable as a bonegraft.

A significant feature of the composite was that it was osteointegrative,i.e., able to bond to bone tissue. No such calcium phosphate layer wasdetected on PLAGA alone, and currently, osteointegration was deemed asignificant factor in facilitating the chemical fixation of abiomaterial to bone tissue. A second feature of the scaffold was thatthe addition of bioactive glass granules to the PLAGA matrix results ina structure with a higher compressive modulus than PLAGA alone.

The compressive properties of the composite approach those of trabecularbone. In addition to being bioactive, the PLAGA-BG lends greaterfunctionality in vivo compared to the PLAGA matrix alone. Moreover, thecombination of the two phases serves to neutralize both the acidicbyproducts produced during polymer degradation and the alkalinity due tothe formation of the calcium phosphate layer. The composite supports thegrowth and differentiation of human osteoblast-like cells in vitro.

The polymer-bioactive glass composite developed for the experiments wasa novel, three-dimensional, polymer-bioactive biodegradable andosteointegrative glass composite scaffold. The morphology, porosity andmechanical properties of the PLAGA-BG construct have been characterized.BG particle reinforcement of the PLAGA structure resulted in anapproximately two-fold increase in compressive modulus (p<0.05).PLAGA-BG scaffold formed a surface Ca-P layer when immersed in anelectrolyte solution (see FIG. 12A), and a surface Ca-P layer wasformed. No such layer was detected on PLAGA controls. EDXA spectraconfirmed the presence of Ca and P (see FIG. 12B) on the surface. TheCa, P peaks were not evident in the spectra of PLAGA controls.

In vitro formation of a surface Ca-P layer indicates PLAGA-BGcomposite's osteointegrative potential in vivo. The growth anddifferentiation of human osteoblast-like cells on the PLAGA-BG scaffoldswere also examined. The composite promoted osteoblast-like morphologyand stained positive for alkaline phosphatase, and promoted synthesis toa greater extent of Type I collagen synthesis than tissue culturepolystyrene controls.

The porous, interconnected network of the scaffold was maintained after3 weeks of culture (see FIG. 13). Mercury porosimetry (MicromeriticsAutopore Micromeritics, Norcross, Ga.) was used to quantify theporosity, average pore diameter and total surface area of the compositeconstruct. The construct porosity was determined by measuring the volumeof mercury infused into the structure during analysis. In addition, theconstruct (n=6) was tested under compression. BG particle reinforcementof the PLAGA structure resulted in approximately two-fold increase incompressive modulus (see Table II, p<0.05).

TABLE II Pore Elastic Compressive Scaffold Average Diameter ModulusStrength Type Porosity (μm) (MPa) (MPa) PLAGA 31% 116 26.48 ± 3.47 0.53± 0.07 PLAGA-BG 43% 89 51.34 ± 6.08 0.42 ± 0.05

Porosity, pore diameter, and mechanical properties of the scaffold maybe variable as a function of microsphere diameter and BG content. Thegrowth and differentiation of human osteoblast-like cells on thePLAGA-BG scaffolds were also examined. The composite supportedosteoblast-like morphology and stained positive for alkalinephosphatase.

The porous, interconnected network of the scaffold was maintained after3 weeks of culture (see FIG. 13). The synthesis of type I collagen wasfound to be the highest on the composite, as compared to the PLAGA andtissue culture polystyrene (TCPS) controls (n=3, p<0.05) (see FIG. 14).

The effects of bovine osteoblast and fibroblast co-culture on theirindividual phenotypes were examined. The cells were isolated usingprimary explant culture. The co-culture was established by firstdividing the surfaces of each well in a multi-well plate into threeparallel sections using sterile agarose inserts. ACL cells andosteoblasts were seeded on the left and right surfaces respectively,with the middle section left empty. Cells were seeded at 50,000cells/section and left to attach for 30 minutes prior to rinsing withPBS. The agarose inserts were removed at day 7, and cell migration intothe interface was monitored. Control groups were fibroblasts alone andosteoblasts alone.

In time, both ACL fibroblasts and osteoblasts proliferated and expandedbeyond the initial seeding areas. These cells continued to grow into theinterfacial zone, and a contiguous, confluent culture was observed. Allthree cultures expressed type I collagen over time. The co-culture groupexpressed type II collagen at day 14, while the control fibroblast didnot. Type X collagen was not expressed in these cultures, likely due tothe low concentration of b-GP used. Alizarin Red S stain intensity wasthe highest for the osteoblast control, (see FIG. 15C) followed by theco-cultured group (see FIG. 15B). Positive ALP staining was alsoobserved for osteoblast control and co-culture groups (see FIGS. 16C and16B, respectively).

Scaffold of four continuous, graded layers with different sizes ofmicrospheres was formulated (see FIGS. 17A-17F). Layered inhomogeneitywas pre-designed into the scaffold. Due to differences in packingefficiency between different sizes of microspheres, the porosity of thescaffold decreases from layers of large microsphere to those consistingof small microspheres. PLAGA-BG composite microspheres were produced viathe emulsion method. Three layers of PLAGA-BG microspheres of differentdiameters (250-300, 300-355, 355-500 μm, from top to bottom) were used,shown in FIGS. 17A-17F. Microsphere layers were sintered at 70° C. for20 hours.

Image analysis confirmed that pore size increased from bottom to top ofscaffold. For the growth of ACL fibroblasts on the scaffold, anothertype of multi-phased scaffold was fabricated using a PLAGA mesh(Ethicon, N.J.) and two layers of PLAGA-BG microspheres. The layers weresintered in three stages in a Teflon mold. First the mesh was cut intosmall pieces and sintered in the mold for more than 20 hours at 55° C. Alayer of PLAGA-BG microspheres with diameter of 425-500 μm was thenadded to the mold. This layer was sintered for more than 20 hours at 75°C. The final layer consisted of PLAGA-BG microspheres with diametergreater than 300 μm. The scaffolds and three distinct regions werereadily observed (see FIGS. 18A-18C).

Kinetics of Ca-P layer formation on BG surfaces was related to changesin surface zeta potential in a simulated body fluid (SBF). The chemicaland structural changes in BG surface Ca-P layer were characterized usingFourier transform infrared spectroscopy (FTIR), SEM and energydispersive x-ray analysis (EDXA). FTIR provides information on thedegree of crystallinity (amorphous vs. crystalline) of the Ca-P layerformed (see FIG. 6) as well as the functional groups present on BGsurface (carbonated Ca-P layer versus non-carbonated, proteinadsorption, etc.). FTIR is much more surface sensitive than X-raydiffraction in detecting the Ca-P crystalline structures when thesurface layer is only several microns in thickness. SEM combined withEDXA is a powerful tool in relating elemental composition to specificsurface morphology and distributions (see FIGS. 7B and 7C). EDXAprovides a direct calculation of Ca/P ratio (Ca/P=1.67 for bone mineraland crystalline Ca-P layer) when appropriate standards are used. FTIR,SEM, and EDXA are complimentary techniques which together providequantitative data on the crystallinity, composition of and functionalgroups pertaining to the Ca-P layer.

Evaluation of the effects of co-culturing on the growth and phenotypicexpression of osteoblasts and chondrocytes. Osteoblasts were seededdirectly on high density chondrocyte micromasses. Specific effects ofco-culture on the expression of chondrogenic markers were observedprimarily at the top surface interaction zone instead of within themicromass. Alcian blue staining (see FIG. 19B) revealed characteristicperi-cellular sulfated GAG deposition by chondrocytes. GAG depositionwas found largely within the micromass, instead of at the co-culturezone where elongated osteoblasts and chondrocytes were located. SulfatedGAG was not detected in the predominantly osteoblast monolayersurrounding the micromass. Surface chondrocytes may havede-differentiated due to co-culturing with osteoblasts. The expressionof type I collagen was observed to be distributed mainly on the topsurface of the co-cultured mass (FIG. 19C), where osteoblasts werelocated. Type I was also found at the primarily osteoblastic monolayersurrounding the micromass (see FIG. 19C, left). No type I collagenexpression was observed in the chondrocyte-dominated center and bottomsurface of the micromass. High expression of type II collagen wasobserved within the micromass (see FIG. 19D).

As types I and II collagen were detected at the surface, it is possiblethat due to co-culture, chondrocytes and osteoblasts were forming anosteochondral-like interface at the surface interaction zone. AlizarinRed (ALZ) staining revealed that there was limited mineralization in theco-cultured group, while the osteoblast control stained increasinglypositive for calcium. It is likely that co-culture with chondrocytes mayhave delayed osteoblast mineralization. Preliminary PCR results (seeFIGS. 20A and 20B) showed that the 7 day co-culture group expressedtypes II and X collagen, as detected by RT-PCR.

Effects of media additives on the growth and mineralization ofosteoblasts and human ACL fibroblasts (hACL) were examined. Duringmineralization, ALP reacted with β-glycerophosphate (β GP) and thephosphate product was utilized for mineralization. Concentrations (0,1.0, 3.0, 5.0 mM) effects were examined over time. No significant changein cell number was observed for the [βGP] investigated. At 1.0 mM, asignificant difference between 1-day & 7-day samples (p<0.05) wasobserved. No differences were found between 1.0 mM and 3.0 mM cultures.ALZ stains for the osteoblast cultures were more intense for 3.0 mM thanfor 1.0 mM. Ectopic mineralization was observed for hACL cultures at 3.0mM suggesting a potential change in cell phenotype.

Interaction of osteoblasts and chondrocytes on a 3-D composite scaffoldduring co-culture was examined. Scaffolds seeded with only osteoblastsor chondrocytes at the same densities served as controls. Bothshort-term and long-term co-culture experiments were conducted.Extensive SEM analysis revealed that significant interactions occurredbetween osteoblasts and chondrocytes during co-culture. Differences incellular attachment were observed between the chondrocyte controlscaffolds and the co-cultured scaffolds. On the co-cultured scaffolds,focal adhesions were evident between the spherical chondrocytes and thesurface, indicated by the arrow in FIG. 21B.

No comparable focal adhesions were observed on the chondrocyte controlsat the same time point. Chondrocyte morphology changed over time as itassumed a spherical morphology in the first 8 hours, and then spread onthe surface of the microspheres (see FIG. 22A). The nodules on thesurface of the microspheres correspond to the flattened chondrocytes.These nodules were likely chondrocytes instead of calcium phosphatenodules, since calcium phosphate nodules were approximately 1-5 μm indiameter at the culture duration observed and these nodules were ˜10 μm,approximately the diameter of an ovoid cell. After 7 days of culture,the co-culture group exhibited extensive matrix production (see FIG.22C) and expansion on the scaffold.

Examination of the ACL-bone interface confirmed existence of a mineralgradient across the insertion zone and correlation to changes inmaterial properties. Multi-phased scaffolds with controlled morphologyand porosity were fabricated. The osteochondral graft developed fromco-culture on PLAGA-BG and hydrogel scaffold supported growth ofmultiple matrix zones with varied GAG and mineral content. BMSCsdifferentiated into ligament fibroblast and produced a functionalextracellular matrix when cultured with growth factors on a fiber-basedscaffold. Mineral content, distribution, and chemistry at the interfaceand on the scaffold were quantifiable using a complimentary set ofsurface analysis techniques (FTIR, SEM, EDAX, μCT). Electron microscopyexamination of the ACL-bone interface revealed insertion zone includingthree different regions: ligament, fibrocartilage-like zone, and bone.Co-culture of osteoblasts and ligament fibroblasts on 2-D and 3-Dscaffolds resulted in changes in cell morphology and phenotype. Type Xcollagen, an interfacial zone marker, was expressed during co-culture.Multi-phased scaffold with layered morphology and inhomogenousproperties were designed and fabricated. FTIR, SEM and EDXA arecomplimentary techniques which collectively provided qualitative andquantitative information on the Ca-P layer and composition of thecalcium phosphate surface.

Second Set of Experiments

A multi-phased scaffold system with inhomogenous properties (FIG. 23)was designed and evaluated for its ability to support the growth anddifferentiation of multiple cell types. Effects of osteoblast-ligamentfibroblast co-culture on a development of interfacial zone specificmarkers (proteoglycan, types II & X collagen) on the scaffold wereexamined.

The contiguous scaffold included three sequential phases (A-C), withPhase A (polymer fiber mesh with no Ca-P) intended for ligament culture,and Phase C (polymer-ceramic composite with high Ca-P) for boneformation. Phase B (polymer-ceramic composite, lower Ca-P than Phase C),the intermediate region, was where an interfacial zone developed due tothe interaction of these two cell types. The scaffolds were fabricatedfrom PLAGA 50:50, and the same polymer was used throughout. The threephases were sintered together past a polymer glass transitiontemperature to form a multi-phased scaffold. The aspect ratio betweenthe phases of the sintered cylindrical scaffold was as follows:A:B:C=2:1:2, and the as-made, complete construct was 1.0 cm in lengthand 0.40 cm in diameter (see FIG. 23).

The mineral gradient was created by incorporating differentconcentrations of bioactive glass (BG) particles during the microspheresynthesis process. BG wt % was correlated to the Ca-P content of theinterface by comparative EDXA analysis of the Ca-P surface developedthrough immersion in a simulated body fluid following awell-characterized method to create Ca-P layer on bioactive glasssurfaces as described by Lu et al. (2000) and incorporated by referenceherein.

When a specific BG wt % was correlated with the Ca-P distribution andCa/P ratio of either the bone or the cartilage region as describedabove, scaffolds were fabricated based on this wt %.

The three phases of the scaffold were inhomogeneous in properties, withzonal differences in mineral content and matrix morphology (see TableIII).

TABLE III Phase A Phase B Phase C (Ligament) (Interface) (Bone)Composition PLAGA PLAGA PLAGA 50:50 50:50/BG 50:50/BG (no BG) (lower)(higher) Porosity/Pore 40%, 40%, 40%, Diameter 100 μm 100 μm 100 μmMatrix Fiber Mesh Microsphere Microsphere Morphology Based Based

The differences mimic the ACL-bone interface and facilitate the growthof different tissues. Phase C has a high mineral content compared toPhase A. While the three phases share the same polymer composition, theydiffer in weight % of BG. A positive correlation exists between scaffoldstiffness and mineral content of the phase.

The three phases also differ in morphology, with Phase A composed of aporous fibrous mesh, and Phases B and C made of microsphere-based porousscaffold. Post-fabrication characterization of the scaffold includedporosity, average pore size, total surface area, as well as mechanicalproperties under compression. Scaffold porosity was held constant at 40%with a pore diameter of 100 μm, with focus on the effect of mineralcontent on cellular response as a more relevant parameter in controllingfibroblast phenotype or dedifferentiation into chondrocytes. Growth anddifferentiation of osteoblasts and ligament fibroblasts co-cultured onthe scaffold were examined. Osteoblasts were seeded on Phase C whileligament fibroblasts were seeded on Phase A.

The growth and differentiation of cells on the scaffold was monitored asa function of culturing time (1, 3, 7, 14, 21 days). Cell proliferation,ligament phenotypic expression (fibronectin, type I, III, II collagensynthesis, laminin, fibronectin) and osteoblast phenotype (alkalinephosphatase, type I collagen, osteocalcin, mineralization) wereexamined. Expression of interface-specific markers such asproteoglycans, types II and X collagen were determined to assess changesin fibroblast phenotype.

The three phases of the scaffold differed in composition and morphology,while the same porosity and pore diameter were maintained. Focus wasplaced on the mineral content of the scaffold for two reasons: 1) it isa more relevant parameter for consideration of the varied mineraldistribution within the ACL-bone interface; and 2) mineral content wasutilized to direct fibroblast phenotype change or dedifferentiation intochondrocytes.

A component of the polymer ceramic composite scaffold was polylactide(PLA) which degrades via hydrolysis into lactic acid, which maycontribute to changes in ligament fibroblast phenotype. Increasedmineralization by ligament fibroblasts was observed with increasingconcentration of β-glycerophosphate, a media additive commonly used inosteoblast cultures.

The effects of co-culture were evaluated in conjunction with scaffoldmineral content. Multiple cell types were considered because theinsertion site was made up of four zones, each dominated by a specificcell type. Cell to cell interactions played a significant role indictating the formation of the interface between ligament and bone.Examination of osteoblast and ligament fibroblast co-culturesestablished that both cell types proliferated and expanded beyond theinitial seeding areas, and that a contiguous and confluent culture wasobserved at the interface after two weeks. Preliminary studies revealedthat co-culture and/or interactions with chondrocytes may have delayedosteoblast-mediated mineralization. Type X collagen was found in theosteoblast-chondrocyte co-cultured samples.

Third Set of Experiments

An objective of the experiments (described below) was to develop athree-dimensional (3-D), porous composite of polylactide-co-glycolide(PLAGA) and 45S5 bioactive glass (BG) that is biodegradable, bioactive,and suitable as a scaffold for bone tissue engineering (PLAGA-BGcomposite). Additional objectives of the study were to examine themechanical properties of a PLAGA-BG matrix, evaluate the response ofhuman osteoblast-like cells to the PLAGA-BG composite, and evaluate theability of the composite to form a surface calcium phosphate layer invitro. Structural and mechanical properties of PLAGA-BG were measured,and the formation of a surface calcium phosphate layer was evaluated bysurface analysis methods. The growth and differentiation of humanosteoblast-like cells on PLAGA-BG were also examined. The addition ofbioactive glass granules to the PLAGA matrix resulted in a structurewith higher compressive modulus than PLAGA alone. Moreover, the PLAGA-BAcomposite was found to be a bioactive material, as it formed surfacecalcium phosphate deposits in a simulated body fluid (SBF), and in thepresence of cells and serum proteins. The composite supportedosteoblast-like morphology, stained positively for alkaline phosphatase,and supported higher levels of Type I collagen synthesis than tissueculture polystyrene controls. A degradable, porous, polymer bioactiveglass composite possessing improved mechanical properties andosteointegrative potential compared to degradable polymers ofpoly(lactic acid-glycolic acid) alone was successfully developed.

Polylactide-co-glycolide 50:50 co-polymer (PLAGA, Mw50,000, AmericanCyanamide, Sunnyvale, Calif.) and 4555 bioactive glass (BG, MO-SCICorporation, Rolla, Mo.) granules were used to fabricate the composite(PLAGA-BG) discs and microspheres. FIG. 24 is a schematic of thesynthesis process of some forms of PLAGA-BG composite used in thisstudy. Specifically, PLAGA-BG discs were formed through the traditionalsolvent-casting process, where PLAGA and BG granules were first mixedaccording to a polymer to ceramic weight ratio of 1:3 and dissolved inmethylene chloride. The solution was then slowly poured into a Teflonmold and allowed to cool overnight in a −20° C. freezer. The resultantpolymer-ceramic film was bored into 1-cm wide and 0.1-mm thick discs.The discs were then dried overnight to remove any residual solvent(Lyph-lock 12, Labconco, Kansas City, Kans.).

PLAGA-BG composite microspheres were formed through a water-oil-wateremulsion. Specifically, PLAGA granules were first dissolved in methylenechloride, and BG particles (<40 μm) were added to achieve a 25% mixture.The mixture was then poured into a 1% polyvinyl alcohol (Polysciences,Warrington, Pa.) solution. The suspension was stirred constantly, andthe spheres were allowed to harden in the polyvinyl alcohol solution.The resultant microspheres were then washed, vacuum filtered, and driedat room temperature. Next, the composite microspheres were sifted usinga mechanical sifter to a final size range of 100-200 μm. The cylindricalconstruct, averaging 0.5 cm in width and 1.0 cm in height, wasfabricated by heating the microspheres at 70° C. for 20 h in astainless-steel mold.

Before in vitro evaluations, the morphology, porosity and mechanicalproperties of the PLAGA-BG construct were determined. Poreinterconnectivity, morphology, and the bonding of microspheres withinthe construct was examined by scanning electron microscopy (SEM, Amray1830-D4), at an acceleration voltage of 20 keV. Elemental composition ofthe composite surface was determined by energy-dispersive X-ray analysis(EDXA). Mercury porosimetry (Micromeritics Autopore III, Micromeritics,Norcross, Ga.) was used to measure the porosity, average pore diameter,and total surface area of the composite construct. In this method, theconstruct porosity was determined by measuring the volume of mercuryinfused into the structure during analysis. In addition, the construct(n=6) was tested under compression using the Instron ServohydrolicSystem 8500 (Instron, Canton, Mass.), with a ramp speed of 0.02 cm/s.The compressive strength and elastic modulus of the construct weredetermined. PLAGA scaffolds without BG served as controls.

The composite discs were immersed for 1, 7, and 14 days in a simulatedbody fluid (SBF) whose ion concentration is similar to that ofextracellular fluid. PLAGA discs without BG served as controls. Asurface area to volume ratio of 1.0 cm-1 was maintained for allimmersions. The pH of the solution as a function of immersion time wasmeasured. Perfect sink conditions were maintained during the immersionstudy. SEM (Amray 1830-D4) and EDXA were used to monitor the formationof a Ca-P layer on composite films.

Human osteosarcoma cells (SaOS-2) were cultured in Medium 199 (M199,Sigma Chemicals, St. Louis, Mo.), supplemented with 10% fetal bovineserum (Life Technologies, Rockville, Md.), L-glutamine, and antibiotics.The cells were grown to confluence at 37° C. and 5% CO2. Under theseconditions, the osteoblastic phenotype of SaOS-2 was maintained for upto at least four weeks of culture, with positive expression of alkalinephosphatase, type I collagen, osteocalcin, and formation of mineralizedcultures.

SaOS-2 cells were seeded on the porous, PLAGA-BG scaffolds (n=3) at thedensity of 5×104 cells/cm2, and were cultured in 12-well plates (FisherScientific, Fair Lawns, N.J.) for up to 3 weeks. PLAGA alone and tissueculture polystyrene (TCPS) served as control groups. Once the cells havegrown to confluence, at two weeks from the start of culture,mineralization medium containing 3.0 mM of β-glycerophosphate and 10μg/ml of L-ascorbic acid were added to the culture.

Cell adhesion and growth morphology on the 3-D construct were monitoredusing SEM (20 keV). Alkaline phosphatase staining was performed at eachculturing time point, using a standard histochemical assay. The sampleswere incubated for 30 min with Napthol AS-Bi (Sigma), phosphate salt,N,N-dimethyl formamide (Sigma), and Fast Red (Sigma) at 37° C. Thesamples were then fixed in 2% paraformaldehyde for 30 min at 4° C. Thesynthesis of type I collagen by SaOS-2 cells was quantified using amodified ELISA.

The formation of mineralized nodules was examined by SEM-EDXA.Mineralization was further ascertained using Alizarin Red S staining forcalcium. Briefly, the samples were washed with deionized H2O, fixed with2% paraformaldehyde and incubated in 2% Alizarin Red S solution for 5min. The samples were then washed with deionized water and viewed underthe microscope.

Data in the graphs are presented in the form of mean±standard deviation(mean±SD), with n equal to the number of samples analyzed per immersiontreatment. One-way analyses of variance (ANOVA) and the Student's t-testwere used to compare the mechanical testing data (n=6), porosimetryresults (n=3), as well as the collagen synthesis data (n=3). Statisticalsignificance was evaluated at the p<0.05.

SEM examination of the PLAGA-BG discs revealed a homogenous distributionof the BG particles within the PLAGA phase. In addition, the compositesin disc form as well as microsphere form were visually more opaque thanPLAGA alone, largely because of the addition of BG. Sintering of themicrospheres resulted in a well-integrated structure, with themicrospheres joined at the contact necks. SEM analysis revealed that a3-D, interconnected porous network was found throughout the compositeconstruct. Elemental analysis using EDXA showed that the compositesurface was largely made up of C, Na, Si, Ca, and P before anyimmersions.

FIG. 25 shows a table which summarizes the result from structuralcharacterizations of the as-fabricated composite scaffold. BGparticle-reinforcement of the PLAGA structure resulted in a neartwo-fold increase in compressive modulus. The structural and mechanicalproperties of the scaffold can be systematically optimized by varyingmicrosphere and scaffold fabrication parameters. Porosimetry analysisrevealed that the 3-D composite measured an average porosity of 43%,with a mean pore diameter of 89 μm. The PLAGA control scaffold exhibited31% total porosity and a mean pore diameter of 116 μm. The PLAGA-BGcomposite possessed a higher elastic modulus (51.336±6.080 MPa versus26.479±3.468 MPa) than the control PLAGA scaffold. Although the meanswere different, the compressive strength of the composite at 0.417±0.054MPa was not statistically different from that of the PLAGA control(0.533±0.068 MPa), at p<0.05.

The bioactivity of the composite was determined by monitoring theformation of a calcium phosphate layer on the composite discs in a SBF.The composite was found to be bioactive because it formed a calciumphosphate layer on its surface after immersion in SBF for 7 days.SEM-EDXA results showed that an amorphous calcium-phosphate layer wasfound on the composite surface after 7 days of immersion, whereas nosuch layer was detected on the control polymer without bioactive glassparticles for the same duration. In particular, polymer-ceramiccomposite (PLAGA-BG) which were immersed in simulated body fluid (SBF)for 14 days formed a surface calcium phosphate layer (Ca, P presenceconfirmed by X-ray analysis as summarized in FIG. 26). No such layer wasfound on the PLAGA control without 45S5 bioactive glass. The composite(PLAGA-BG) surface was covered with calcium phosphate nodules after 14days of immersion. In contrast, the PLAGA control surface, afterimmersion for 14 days in SBF, did not form a calcium phosphate layer,but began to exhibit surface pores formed due to the degradation of thepolymer.

FIG. 26 shows EDXA spectra of the PLAGA-BG composite immersed in a SBFfor 14 days. The composite surface still contained C, Si, Ca, and P,whereas the Cl peak was detected after immersion in SBF. A surfacecalcium phosphate layer has formed on the PLAGA-BG composite surface.The Ca and P peaks were not found in the spectra of PLAGA controls.

The microsphere-based, porous, PLAGA-BG composite supported the growthand phenotypic expression of human osteoblast-like cells. Media pHvariation was measured for the full duration (3 weeks) of cell culturewith PLAGA-BG and PLAGA, and physiological pH (7.3-7.7) was maintainedin all cultures for up to 3 weeks. There was no significant change insolution pH after 2 weeks of culture with osteoblast-like cells, andculture media was exchanged every other day to remove metabolic productsand supply fresh nutrients to the cells. Extensive cellular growth wasdetected on the scaffold surface as well as within the PLAGA-BGcomposite. In addition, the porous network of the scaffold wasmaintained even after 3 weeks of culture. In many areas, cellular growthhad bridged two or more microspheres while maintaining the porousstructure. SEM analysis revealed the synthesis of collagen-like fibersby the SaOS-2 cells. All cultures stained positively for the synthesisof alkaline phosphatase, although a much higher intensity of stain wasobserved in cultures with the PLAGA-BG scaffold than for PLAGA cultures.

As shown in FIG. 27, the synthesis of type I collagen by SaOS-2 cellsincreased with culturing time, with the highest amount found on PLAGA-BGcomposite (0.146±0.006 μg), as compared to PLAGA (0.132±0.006 μg), andTCPS controls (0.073±0.005 μg). The expression of type I collagen bySaOS-2 cells cultured on the composite was significantly higher thancells grown on TCPS controls, (p<0.05). There was a trend towards higherType I collagen synthesis on the PLAGA-BG composite compared to PLAGAalone, but this was not found to be significant. (p=0.06) The formationof a mineralized matrix was confirmed by positive staining with AlizarinRed S and elemental analysis in which Ca and P were detected on PLAGABGscaffolds cultured with SaOS-2 cells. Alizarin stain intensity increasedwith culturing time. The mineralized nodules were not observed on PLAGAor TCPS controls after 2 weeks of culture, before the addition of themineralization medium. After 1 week of culturing with the mineralizationmedium, mineralization as reflected in staining intensity, was much lesson the control substrates than on PLAGA-BG.

SEM and EDXA analyses confirmed the formation of calcium phosphatenodules on the composite surface after only 3 days of culture, beforethe addition of the mineralization medium. These calcium phosphatenodules are similar in size and shape as observed on PLAGA-BG discs inthe SBF. In time, the Ca-P nodules increased in size and formed largeraggregates, indicating that the PLAGABG composite was bioactive invitro. The relative Ca to P peak ratio of the deposits decreased as afunction of culturing time. These results collectively suggest that thecomposite was bioactive, and was capable of forming a surface calciumphosphate layer.

Fourth Set of Experiments

The degree of graft integration is a significant factor governingclinical success and it is believed that interface regenerationsignificantly improves the long term outcome. The approach of this setof experiments was to regenerate the ACL-bone interface throughbiomimetic scaffold design and the co-culture of osteoblasts andfibroblasts. The interface exhibits varying cellular, chemical, andmechanical properties across the tissue zones, which can be explored asscaffold design parameters. This study describes the design and testingof a multi-phased, continuous scaffold with controlled heterogeneity forthe formation of multiple tissues. The continuous scaffold consists ofthree phases: Phase A for soft tissue, Phase C for bone, and Phase B forinterface development. Each phase was designed with optimal compositionand geometry suitable for the tissue type to be regenerated. Fibroblastswere seeded on Phase A and osteoblasts were seeded on Phase C, and theinteractions of osteoblasts and fibroblasts (ACL and hamstring tendon)during co-cultures on the scaffolds were examined in vitro.

Phases A, B and C consist of poly(lactide-co-glycolide) (PLAGA, 10:90)woven mesh, PLAGA (85:15) microspheres, and PLAGA (85:15)/BioactiveGlass (45S5,BG) composite microspheres, respectively. The microsphereswere formed via a double emulsion method, and the continuousmulti-phased scaffolds were formed by sintering above the polymer T_(g).Scaffold porosity and pore diameter were determined by porosimetry(Micromeritics, n=3) and the samples were tested under uniaxialcompression (MTS 810, n=5) at 1.3 mm/min up to 5% strain with 10 Npreload.

Bovine and human osteoblasts (bOB and hOB), and bovine ACL fibroblasts(bFB) and human hamstring tendon fibroblasts (hFB) were obtained throughexplant culture. In experiment I, bOB and bFB (5×10⁵ cellseach/scaffold) were co-cultured on the scaffold, and cell viability,attachment, migration and growth were evaluated by electron andfluorescence microscopy. The bOB were pre-labeled with CM-DiI, and bothcell types were labeled with calcein AM (Molecular Probes) prior toimaging. Matrix production and mineralization were determined byhistology. After ascertaining cell viability on the scaffolds, a moreextensive experiment using hOB and hFB was conducted in which cellproliferation and differentiation and above analyses were investigated.The mechanical properties of the seeded scaffolds were also measured asa function of culture time.

Compression testing of scaffolds indicated an average modulus of 120±20MPa and yield strength of 2.3 MPa. The intrusion volume, porosity andpore diameter data are summarized in the table shown in FIG. 28-1.

The fibroblasts and osteoblasts were localized primarily at the two endsof the scaffolds after initial seeding, with few cells found in Phase B.After 28 days, both cell types migrated into Phase B (FIG. 28-2B), andextensive cell growth was observed in Phases A and C (FIGS. 28-2A and28-2C).

Extensive collagen-rich matrix production was found throughout the threephases at day 28 (FIGS. 28-3A and 28-3B).

The biomimetic, multi-phased scaffolds supported the growth and ECMproduction of both osteoblasts and fibroblasts. After 28 days ofculture, collagen production was evident in all three phases andmineralized matrix was found in the bone and interface regions.Osteoblast and fibroblast interaction at the interface (Phase B)suggests that these cells may play a significant role in the developmentof a functional insertion site. These findings demonstrate that thisnovel scaffold is capable of simultaneously supporting the growth ofmultiple cell types and can be used as a model system to regenerate thesoft tissue to bone interface. Additional studies can focus on scaffoldoptimization and the development of the interface on the novel scaffold.

Fifth Set of Experiments

This set of experiments is directed to the development of a multi-phasedscaffold with controlled heterogeneity for interface tissue engineering.This continuous scaffold is comprised of three phases with Phase Adesigned for ligament formation, Phase C for bone, and Phase B forinterface development. The design objective was to formulate a scaffoldthat is able to support the growth and differentiation of bothosteoblasts and ligament fibroblasts. Two design parameters were variedamong the three phases: mineral (Ca/P) content and geometry. This studyintroduces a 3-D biomimetic substrate for interface development. Theinteraction of osteoblasts and ACL fibroblasts during co-culture on themulti-phased scaffold were examined. An objective of the study was todemonstrate that both cell types proliferate and elaborate a collagenlike matrix on the 3-D scaffolds.

Two types of scaffolds were fabricated. The first type is comprisedentirely of microspheres formed via a double emulsion method. Phase Aconsists of poly(lactide-co-glycolide) 50:50 (PLAGA), Phase C ofPLAGA/Bioactive glass (PLAGA-BG) composite microspheres, and Phase Bcontains a mixture of PLAGA and PLAGA-BG. For the second type ofscaffold which has a different geometry and degradation rate, Phase Aconsists of PLAGA (10:90) woven mesh, Phase C of PLAGA 85:15/BGmicrospheres, and Phase B contains PLAGA (85:15) microspheres. Thecontinuous multi-phased scaffolds were formed by sintering above theglass transition temperature.

Bovine osteoblasts and ACL fibroblasts were obtained from explantcultures of tissue isolated from neonatal calves. The cells werecultured in Dulbecco's Modified Eagles Medium (DMEM, Mediatech),supplemented with 10% fetal bovine serum, L-glutamine, and 1%penicillin/streptomycin (Mediatech).

Scaffolds were sterilized by ethylene oxide and fibroblasts were seededat a density of 5×10⁵ cell/scaffold onto Phase A, while osteoblasts wereseeded at 5×10⁵ cell/scaffold on Phase C. Phase B was left unseeded andthe migration of osteoblasts and fibroblasts into this interfacialregion was examined. The osteoblasts were labeled with CM-DiI celltracer (Molecular Probes), and their location was tracked with respectto fibroblasts and each phase of the scaffold. The scaffolds werecultured in supplemented DMEM for up to 28 days. Ascorbic acid (10μg/mL) and 3 mM β-glycerophosphate were added to the cultures at day 7.

Cell migration, attachment and growth were examined using scanningelectron microscopy (5 kV, JEOL 5600LV). Cell viability and migrationwere evaluated by fluorescence microscopy (Zeiss Axiovert 40) usingcalcein AM tracer (Molecular Probes). Matrix production andmineralization were determined via histology. The samples were fixed,embedded and sectioned, after which Trichrome, von Kossa and PicrosiriusRed stains were performed.

At day 0, SEM analysis showed that a large number of cells attached toPhase A and C of the scaffolds (FIG. 29-1A). Fluorescence microscopyrevealed that fibroblasts and osteoblasts were localized primarily atopposite ends of the scaffolds after initial seeding, with very fewcells found in Phase B (FIGS. 29-2A through 29-2C). At day 28, SEManalysis revealed that both cell types elaborated extracellular matrix(ECM) on Phases A and C (FIGS. 29-1B and 29-1C) with some matrixformation observed in Phase B (FIG. 29-1D). Fibroblasts were foundlargely in Phase A and osteoblasts in Phase C (FIGS. 29-2D and 29-2F),with a mixture of cell types found in Phase B (FIG. 29-2E).

Histological analyses confirmed cell migration into Phase B and matrixproduction throughout the three phases of the scaffold at day 28 (FIGS.29-3A1 through 29-3A3). The collagen-rich matrix (FIGS. 29-3B1 and29-3B2) seen in all three phases and osteoblast-mediated mineralizationwere observed on the surface of the PLAGA-BG microspheres (FIG. 29-3C,see arrow).

The biomimetic, multi-phased scaffolds supported the growth and ECMproduction by both osteoblasts and fibroblasts. After 28 days ofculture, collagen production was evident in all three phases andmineralized matrix was found in the bone and interface regions only.Osteoblast and fibroblast interaction at the interface (Phase B)suggests that these cells may serve a significant role in thedevelopment of a functional insertion site. The results demonstrate thatthis novel scaffold is capable of simultaneously supporting the growthof multiple matrix zones. Additional studies can examine the effects ofcell-cell interactions at the interface region and optimize the scaffoldfor clinical utilization.

Sixth Set of Experiments

It is believed that fibroblasts and osteoblasts interactions play asignificant role in interface formation. In vivo, fibroblasts andosteoblasts form a fibrocartilage layer within the bone tunnel. Sincethe natural interface spans less than 400 μm, a novel micro-co-culturemodel was developed that utilizes microfluidics to exert spatial controlin cell distribution. This can be used to determine how cell-cellinteractions may regulate interface remodeling locally at themicro-scale. The fabrication parameters of this model were optimized andinitial osteoblastic and fibroblastic responses were examined.

Channels were designed having a bimodal non-intersecting serpentinegeometry with 200 μm features. The design was implemented on siliconwafers using SU-8 25 (Microchem) photoresist and a mold patterned usingPolydimethylsiloxane (PDMS, Dupont). In this design, osteoblast andfibroblast channels were first separated by PDMS, which was laterremoved to allow cell to cell interactions.

In order to optimize the channel depth for subsequent co-culturestudies, the spin-coating durations (30, 45, 60 and 90s) were varied.Cell seeding time was optimized by incubating the cells within thechannels for 1, 3, 6, and 24 hours prior to removal of the PDMS followedby live-dead staining.

Bovine primary osteoblasts and fibroblasts were obtained from explantcultures. The cells were grown in supplemented DMEM (10% FBS, 1% NEAAand 1% antibiotics) at 37° C. and 5% CO₂. Osteoblast or fibroblastsuspension (20×10⁶ cells/ml) was perfused into its respectivemicrochannels. Cells were allowed to attach for 1 hr prior to PDMSremoval. Cell migration was tracked by labeling fibroblasts with CM-DiIand osteoblasts with CFDA-SE (Molecular Probes) prior to seeding.

Analyses were performed at days 1, 2, and 6 following PDMS removal.Alkaline Phosphatase (ALP) activity was ascertained with fast-blue stain(Sigma), while type-I collagen deposition was examined byimmunohistochemistry.

A spin-coating duration of 30 s was chosen to balance channel depth anduniformity. Based on the cell viability, the optimal cell attachmenttime within the channels was 1 hr (FIG. 30-2 a). Both cell typesmigrated and proliferated beyond their initial seeding zone (FIGS. 30-1a through 30-1 d) and grew into physical contact by day 1 (FIGS. 30-1 eand 30-1 f). Local confluency and cross-migration were observed at day2. ALP activity was observed in the osteoblast region (FIG. 30-2 b),while type-I collagen was found in all regions (FIG. 30-2 c).

A successful micro-co-culture model was developed and initialexamination of the interactions between osteoblasts and fibroblasts in amicro-co-culturing environment was performed. Cells proliferated beyondthe initial seeding region and maintained their phenotypes as indicatedby ALP activity of osteoblasts and type-I collagen deposition of bothcell types. The cell-to-cell cross-migration at day 2 offered a host ofhomotypic and heterotypic cell interactions. Micropatterning of cellsoffers an unique opportunity to control the local micro-environment andpermit the in-depth examination of cell-cell interactions. Thisunderstanding can aid in the identification of mechanisms drivinginterface formation.

Seventh Set of Experiments

This set of experiments was directed to in vitro evaluations of humanosteoblasts and fibroblasts co-cultured on multi-phased scaffolds. Aschematic of the experimental design for the in vitro study is shown inFIG. 31-1. Phase A (meshe) was seeded with human hamstring tendonfibroblast cell suspension. Phase C was seeded with osteoblasts. Cellinteraction in the interfacial Phase B was monitored over time.Acellular scaffolds served as controls.

Cell proliferation in Phases A, B, and C during 35 days of humanhamstring tendon fibroblast and osteoblast co-culture on multiphasedscaffolds is shown in FIG. 31-2. A general trend of increasing cellnumber was observed in each phase over time. Data demonstrates that allthree phases of the scaffold support cellular viability andproliferation. A higher number of cells were seeded on phase A due toits inherently larger surface area compared to phase C.

Mechanical testing data for multiphased scaffolds seeded with humanhamstring tendon fibroblasts and human osteoblasts over 35 days ofculture (n=4) is graphically shown in FIGS. 31-3A and 31-3B. Scaffoldswere tested in uniaxial compression. Compressive modulus (FIG. 31-3A)and yield strength (FIG. 31-3B) were calculated from the resultingstress-strain curves. Both cell seeded (C) and acellular (AC) scaffoldswere examined at days 0, 7, 21, and 35.

Compared to the acellular controls, the cell seeded scaffolds degradedslower and better maintained their structural integrity over time. Theyield strength of the acellular scaffold decreased over 35 days, whilethe seeded scaffolds maintained its yield strength.

Eighth Set of Experiments

The scaffold designed for this study consisted of three phases and werefabricated in four stages (FIG. 32-1). First, Phase A was formed frompolyglactin 10:90 PLGA mesh sheets (Vicryl VKML, Ethicon). Mesh sheetswere cut into small segments (approximately 5 mm×5 mm) and inserted intocylindrical molds (7.44 mm diameter). Molds were heated to 150° C. for20 hours to sinter the segments together to form a cylindrical meshscaffold. The next phase (Phase B) consisted of 100%85:15—poly(DL-lactide-co-glycolide) (PLAGA, Alkermes Medisorb,M_(W)≈123.6 kDa) microspheres formed by a water/oil/water emulsion.Briefly, 1 g PLAGA was dissolved in 10 mL methylene chloride (EMScience, Gibbstown, N.J.) and poured into a mixing 1% PVA surfactantsolution (Sigma Chemicals, St. Louis, Mo.). Microspheres were mixed for4 hours, recovered by filtration, allowed to dry in a fume hoodovernight, then vacuum desiccated for 24 hours. To form the PLAGAmicrosphere phase, ˜0.075 g microspheres were inserted into the samemolds as used previously, and sintered at 55° C. for 5 hours. The lastphase (Phase C) consisted of composite microspheres formed from an 80:20ratio of PLAGA and 45S5 bioactive glass (BG, MO-SCI Corporation, Rolla,Md.). Again, microspheres were formed by emulsion, except with 0.25 gbioactive glass suspended in a solution of 1 g PLAGA in 10 mL methylenechloride. Microspheres (28-30 mg/scaffold) were sintered in the samemolds at 55° C. for five hours. After all three phases were sinteredseparately, Phases A and B were joined by methylene chloride solventevaporation, and then sintered to Phase C for 10 hours at 55° C. in thesame molds. Subsequently, scaffolds were sterilized with ethylene oxide.Final scaffold dimensions are detailed in FIGS. 32-4A and 32-4B.

Human osteoblast-like cells and hamstring tendon fibroblasts wereobtained from explant culture of tissue isolated from humerus trabecularbone and hamstring tendon respectively. Trabecular bone was rinsed withPBS, then cultured in Dulbecco's Modified Eagle's Medium (DMEM,Mediatech, Herndon, Va., USA) supplemented with 10% fetal bovine serum,1% non essential amino acids, and 1% penicillin/streptomycin (Mediatech,Herndon, Va.), and incubated at 37° C. in a 5% CO2 incubator to allowfor cell migration. Hamstring tendon obtained from excess tissueutilized for hamstring tendon ACL reconstruction autografts was mincedand cultured in similarly supplemented DMEM. The first migrations ofcells were discarded to obtain a more uniform cell distribution. Secondmigration, passage 2 osteoblast-like cells and second and thirdmigration, passage 5 hamstring tendon fibroblasts were utilized for theco-culture experiment.

Scaffold dimensions were measured prior to cell seeding and before andafter EtO sterilization. Phase thickness was calculated by imageanalysis, while phase diameter was determined using a digital caliper.Scaffold porosity and pore diameter (Phases A and B: n=3; Phase C: n=1)were determined by mercury porosimetry (Micromeritics Autopore III andAutopore IV 9500, Micromeritics, Norcross, Ga.). The porosity data wereutilized to determine cell seeding densities and cell suspension volumesfor Phases A and C, with the volumes calculated such that fibroblastssuspension remains in Phase A and osteoblasts suspension in Phase C.

Hamstring tendon fibroblasts were seeded at a density of 250,000cells/scaffold in a volume of 40.7 μL/scaffold on Phase A (FIG. 32-2).After allowing the fibroblasts to attach to the scaffolds for 20minutes, the scaffolds were rotated upside down so that Phase C facedupwards. Subsequently, 75,000 osteoblast-like cells were seeded perscaffold in a volume of 12.5 μL. After allowing the osteoblasts toattach to the scaffold for 20 minutes, the scaffolds were covered withDMEM supplemented with 10% FBS, 1% NEAA, and 1% penicillin/streptomycin,and incubated at 37° C. and 5% CO₂. Ascorbic acid at a concentration of20 μg/mL was added beginning at day 7. Media was exchanged every twodays. Scaffolds were cultured in 6-well plates and covered with 7 mL ofsupplemented media per scaffold to minimize pH fluctuations due to rapidpoly(glycolic acid) degradation.

Cell attachment, migration, and proliferation on the multi-phasedscaffolds were examined using SEM (5 kV, JEOL 5600LV) at days 7, 21, and35. The scaffolds were fixed with Karnovsky's glutaraldehyde fixative,and stored at 4° C. for 24 hours. The samples were then rinsed withHank's buffered salt solution two times, and serially dehydrated withethanol. Cross-sections of the scaffold phases were mounted on analuminum post and gold-coated prior to analysis.

Extracellular matrix production and mineralization were determined viahistology at day 35. Scaffolds were rinsed two times with roomtemperature PBS. The scaffolds were then covered with 10% neutralbuffered formalin and stored at 4 degrees C. Samples were plasticembedded using a modification of a procedure developed by Erben.

The scaffolds were first suspended in 2% agarose (low gellingtemperature, cell culture grade, Sigma, St. Louis, Mo.), then seriallydehydrated with ethanol and cleared with xylene substitute (Surgipath,Sub-X, Richmond, Ill.). Following dehydration, samples were embedded inpoly(methyl methacrylate) (Polysciences, Inc., Warrington, Pa.) andsectioned into 10 μm slices. The scaffold sections were stained witheither hematoxylin and eosin, von Kossa or Picrosirius Red stains andimaged with light microscopy.

At days 1, 7, 21, and 35, scaffolds were rinsed twice with PBS andsubsequently the three phases were separated. Each phase was then storedin 0.1% Triton-X at −80° C. Cellular proliferation in each phase wasdetermined by means of PicoGreen DNA quantitation assay. In addition,cellular phenotype for mineralization was evaluated using a quantitativealkaline phosphatase (ALP) assay.

At days 0, 7, 21, and 35, seeded and acellular scaffolds were testedunder uniaxial compression (MTS 810, n=4). The crosshead speed was 1.3mm/min, and the scaffolds were compressed up to 35-40% strain. A 10 Npreload was applied prior to testing. The effects of scaffolddegradation and extracellular matrix production on scaffold compressivemodulus were examined.

Mercury porosimetry data for each phase are summarized in the tableshown in FIG. 32-3. Scaffold dimensions are shown in FIGS. 32-4A and32-4B. The thickness of Phase C decreased significantly (p<0.05) due tocontraction during the EtO sterilization (FIG. 32-4A). In addition, thethicknesses of all phases were significantly different from each otherafter sterilization. Scaffold diameters also varied due to contractionduring sintering, in the case of Phase A, and contraction of Phase Cduring sterilization. The diameters of Phases B and C decreasedsignificantly after sterilization, and the diameters of all phases weresignificantly different from each other after sterilization (p<0.05).During the scaffold fabrication process, microspheres are lost betweenweighing and filling the molds. This loss is mainly due to static chargeaccumulation in one or more of the microspheres, weighing paper, ormold, which prevents a small percentage of the microspheres fromentering the molds. PLAGA-BG microspheres for Phase C generallyexperience a 2.1±1.4% loss in mass, while the PLAGA microspheres forPhase B suffer a loss of 4.0±1.8% (FIG. 32-5). Composite microspheresare generally more statically charged than the PLAGA microspheres;however, the stainless steel mold, used more often for the compositemicrospheres, dissipates charge buildup more readily than the PTFE mold,which is used more often for the PLAGA microspheres, possibly explainingwhy there is a significant loss for Phase B (p<0.05). Mesh for Phase Ais not susceptible to this loss.

Compressive modulus and yield strength were obtained for seeded andacellular control scaffolds at days 0, 7, 21, and 35 of culture. A rapiddecrease in compressive modulus was observed following day 0, possiblydue to rapid initial polymer degradation. By day 35, the seededscaffolds exhibited a greater compressive modulus (FIG. 32-6A) and yieldstrength (FIG. 32-6B), possibly due to cellular extracellular matrix andmineralization compensating loss of scaffold strength due to polymerdegradation.

In this experiment, the cell types were switched from bovine ACLfibroblasts and trabecular bone osteoblast-like cells to human hamstringtendon fibroblasts and trabecular bone osteoblasts due to the increasedclinical relevance of these new cell types. This experiment aimed toacquire quantitative data about cell proliferation and migrationthroughout the three phases, as well as cellular alkaline phosphataseactivity in each phase of the scaffold.

Based on the previous experiment performed with bovine cells, it isapparent that the biomimetic, multi-phased scaffolds support the growthand ECM production of both osteoblasts and fibroblasts. After 28 days ofculture, collagen production was evident in all three phases andmineralized matrix was found in the bone and interface regions.Osteoblast and fibroblast interaction at the interface (Phase B)suggests that these cells may play a significant role in the developmentof a functional insertion site. These findings demonstrate that thisnovel scaffold is capable of simultaneously supporting the growth ofmultiple cell types and can be used as a model system to regenerate thesoft tissue to bone interface. Additional studies can focus on scaffoldoptimization and the development of the interface on the novel scaffold.

Ninth Set of Experiments

The objective of the set of experiments was to incorporate electrospunPLAGA meshes into the multi-phased scaffold design, substituting theEthicon mesh phase, and allowing the entire scaffold to be madein-house.

Electrospinning, short for electrostatic spinning, is a relatively newterm that describes a principle first discovered in the first half ofthe 20^(th) century (see, for example, U.S. Pat. Nos. 1,975,504,2,160,962, 2,187,306, 2,323,025 and 2,349,950 to Formhals, the entirecontents of which are incorporated herein by reference). Electrostaticspinning involves the fabrication of fibers by applying a high electricpotential to a polymer solution. The material to be electrospun, ordissolved into a solution in the case of polymers, is loaded into asyringe or spoon, and a high potential is applied between the solutionand a grounded substrate. As the potential is increased, theelectrostatic force applied to the polymer solution overcomes surfacetension, distorting the solution droplet into a Taylor cone from which ajet of solution is ejected toward the grounded plate. The jet splaysinto randomly oriented fibers, assuming that the solution has a highcohesive strength, linked to polymer chain molecular weight, to preventdroplets from forming instead of fibers in a process known aselectrospraying. These fibers have diameters ranging from nanometerscale to greater than 1 μm and are deposited onto the grounded substrateor onto objects inserted into the electric field forming a non-wovenmesh. Mesh characteristics can be customized by altering electrospinningparameters. For example, fiber diameter and morphology can be altered,including the formation of beads along the fibers, by controllingapplied voltage and polymer solution surface tension and viscosity.Also, fiber orientation can be controlled by rotating the groundedsubstrate. This high degree of customizability and ability to use manydifferent materials, such as biodegradable polymers and silks, grantthis fabrication method a high potential in the development of materialsfor biomedical application. Management of fiber diameter allows surfacearea to be controlled, and polymers with different degradation rates canbe combined in various ratios to control fiber degradation, both ofwhich are significant in drug delivery applications. Also, controllingthe orientation of fiber deposition grants a degree of control over cellattachment and migration. Moreover, the ability to electrospin fibermeshes onto non-metal objects placed in the electric field enables thefabrication of multiphasic scaffold systems.

Here, in order to obtain precise parameters for the mesh fibers,including fiber diameter, morphology, and alignment, the effects ofprocessing parameters on fiber characteristics were studied. Avariable-speed rotating drum was designed and constructed to serve as asubstrate for aligned fibers, and rheological experiments were performedon the polymer solutions to determine the effect of polymerconcentration on solution viscosity and the subsequent effect ofsolution viscosity on fiber diameter and morphology.

In addition to determining the speed of each gear, the effect of eachspeed on fiber alignment was determined qualitatively. A 30% v/v PLAGAsolution was prepared with 60% dimethylformamide and 10% ethanol, andthis solution was electrospun onto the rotating drum at each of the fourspeed settings. The resulting meshes were examined by scanning electronmicroscopy (JEOL 5600LV).

The relationship between polymer concentration (Alkermes 85:15 PLAGA)and solution viscosity was determine by means of a rheological study.Three concentrations of polymer were tested—20%, 30%, and 40% v/v—indimethylformamide (DMF) and ethanol. The composition of each solution islisted in the table shown in FIG. 33-1. Solutions were analyzed using anAdvanced Rheometer AR 2000t. There was variability in the viscositymeasurements (n=1) at different strain rates due to the evaporation ofsolvent during testing. The geometry used for the viscosity measurementswas a 25 mm stainless steel disc. A solvent trap was not used since itis not designed to fit with this geometry and a prior trial using thesolvent trap with another geometry resulted in poor results, possiblybecause water from the solvent trap seal interacted with the polymersolution. Additional trials can use a solvent trap to obtain consistentand reliable values for viscosity. For the present study, averages weretaken of the viscosity measurements taken at strain rates tested afterthe equipment had equilibrated. As a result, there are standarddeviations for the viscosity measurements even with an n of 1.

The surface velocity of the rotating drum was seen to increase withincreased pulley positions from gear 1 to gear 4 (see the table shown inFIG. 33-2). The degree of fiber alignment increased with increasing drumvelocity, as seen in the SEMs of each mesh (see FIGS. 33-3A through33-3D).

It was found that (as expected) the degree of fiber orientationincreased with increasing drum rotational velocity. The image wasanalyzed and a histogram of fiber angles was generated against thehorizontal axis of the image at regular interval across the image. Thus,the degree of alignment of the fibers can be quantified. It is desirableto control the degree of fiber alignment in the electrospun meshes sothat the extracellular environment found at the interface can bemimicked. By producing biomimetic scaffolds, it was intended to directcell growth to reproduce the tissue inhomogeneity found at the nativeACL insertions. In addition to controlling the fiber alignment, it isdesirable to control fiber diameter and morphology. It was previouslydetermined that substituting 10% of the DMF in the polymer solutionswith ethanol reduces the surface tension of the solution and results ina significant reduction in the number of beads formed along the fiberswhen electrospinning PLAGA. This effect was also observed by Fong etal., who reduced the number of beads in electrospun polyethylene oxide)(PEO) meshes by the addition of ethanol. Surface tension of the polymersolution acts to form spheres during the electrospinning process. Byreducing the solution surface tension, the formation of spheres is lessfavorable and straighter fibers result. Fong et al. also determined thatthe addition of ethanol increased the viscosity of the PEO:watersolutions, which also favors the formation of straight fibers, andresults in increased fiber diameter. Deitzel et al. also havedemonstrated a relationship between PEO:water solution viscosity andfiber diameter, with fiber diameter increasing with increasing viscosityaccording to a power law. A relationship between solution viscosity andconcentration of polymer can be determined in order to understand howPLAGA:N,N-DMF viscosity affects fiber diameter and morphology. Theeffect of solution viscosity on fiber diameter and morphology can bedetermined by spinning the various solutions and examining the resultingmeshes by SEM. Other variables can affect the fiber parameters. Bychanging the percentage of polymer, the surface tensions of the polymersolutions also change in addition to the viscosity. Therefore, inaddition to testing the viscosities of each solution, the surfacetension of each solution are measured. It is desirable to keep allvariables constant except for viscosity in order to truly determine theeffect of solution viscosity on fiber characteristics. However, theinterrelation of many of the electrospinning parameters complicates theprocess.

A PLAGA mesh was electrospun directly onto a microsphere scaffold. Thisis one way to incorporate the mesh. In addition, the scaffolds can besecured to the drum and aligned fibers electrospin directly onto thescaffolds. However, because of the high rotational velocities, it isdifficult to secure the scaffolds and prevent them from flying off thedrum when it begins rotating. Alternatively, aligned fiber meshes cansimply be spun separately, and then later sintered to the microspherescaffolds. For example, aligned fiber meshes can be electrospun ontoaluminum foil, then wrapped around a rod with multiple mesh sheetssintered together to obtain a hollow cylinder of aligned fibers.

FIG. 33-4A and 33-4B show scanning electron microscopy (SEM) images ofanother embodiment of multi-phased scaffold, with 85:15 PLAGAelectrospun mesh joined with PLAGA:BG composite microspheres.

REFERENCES

-   1. Abate, J. A., Fadale, P. D., Hulstyn, M. J. & Walsh, W. R.,    “Initial fixation strength of polylactic acid interference screws in    anterior cruciate ligament reconstruction,” Arthroscopy 14, 278-284    (1998).-   2. Allum, R. L., “BASK Instructional Lecture 1: graft selection in    anterior cruciate ligament reconstruction,” Knee 8, 69-72 (2001).-   3. Altman, G. H., et al. “Advanced bioreactor with controlled    application of multi-dimensional strain for tissue engineering,”    Journal of Biomechanical Engineering 124, 742-749 (2002).-   4. Altman, G. H., et al., “Silk matrix for tissue engineered    anterior cruciate ligaments,” Biomaterials 23, 4131-4141 (2002).-   5. American Academy of Orthopaedic Surgeons, “Arthoplasty and Total    Joint Replacement Procedures: United States 1990 to 1997,” (Report,    United States) (1997).-   6. Anderson, K. et al., “Augmentation of tendon healing in an    intraarticular bone tunnel with use of a bone growth factor,” Am. J.    Sports Med. 29, 689-698 (2001).-   7. Batycky, R. P., Hanes, J., Langer, R. & Edwards, D. A., “A    theoretical model of erosion and macromolecular drug release from    biodegrading microspheres,” J. Pharmaceutical Sciences 86, 1464-1477    (1997).-   8. Bellincampi, L. D., Closkey, R. F., Prasad, R., Zawadsky, J. P. &    Dunn, M. G., “Viability of fibroblast-seeded ligament analogs after    autogenous implantation,” J. Orthop. Res. 16, 414-420 (1998).-   9. Benjamin, M., Evans, E. J., Rao, R. D., Findlay, J. A. &    Pemberton, D. J., “Quantitative differences in the histology of the    attachment zones of the meniscal horns in the knee joint of man,” J.    Anat. 177, 127-134 (1991).-   10. Berg, E. E., “Autograft bone-patella tendon-bone plug    comminution with loss of ligament fixation and stability,”    Arthroscopy 12, 232-235 (1996).-   11. Beynnon, B. et al., “A sagittal plane model of the knee and    cruciate ligaments with application of a sensitivity analysis,” J.    Biomech. Eng. 118, 227-239 (1996).-   12. Beynnon, B. D. et al., “The effect of functional knee bracing on    the anterior cruciate ligament in the weightbearing and    nonweightbearing knee,” Am. J. Sports Med. 25, 353-359 (1997).-   13. Blickenstaff, K. R., Grana, W. A. & Egle, D., “Analysis of a    semitendinosus autograft in a rabbit model,” Am. J. Sports Med. 25,    554-559 (1997).-   14. Bolton, C. W. & Bruchman, W. C., “The GORE-TEX expanded    polytetrafluoroethylene prosthetic ligament. An in vitro and in vivo    evaluation,” Clin. Orthop. 202-213 (1985).-   15. Borden, M., Attawia, M., Khan, Y. & Laurencin, C. T. Tissue    engineered microsphere-based matrices for bone repair: design and    evaluation. Biomaterials 23, 551-559 (2002).-   16. Bonfield, W., “Composites for bone replacement,” J. Biomed. Eng.    10, 522-526 (1988).-   17. Boskey, A. L. et al., “The mechanism of beta-glycerophosphate    action in mineralizing chick limb-bud mesenchymal cell cultures,” J.    Bone Min. Res. 11, 1694-1702 (1996).-   18. Brand, J., Jr., Weiler, A., Caborn, D. N., Brown, C. H., Jr. &    Johnson, D. L., “Graft fixation in cruciate ligament    reconstruction,” Am. J. Sports Med. 28, 761-774 (2000).-   19. Brody, G. A., Eisinger, M., Arnoczky, S. P. & Warren, R. F., “In    vitro fibroblast seeding of prosthetic anterior cruciate ligaments.    A preliminary study,” Am. J. Sports Med. 16, 203-208 (1988).-   20. Bromage, T. G., Smolyar, I., Doty, S. B., Holton, E. & Zuyev, A.    N., “Bone growth rate and relative mineralization density during    space flight,” Scanning 20, 238-239 (1998).-   21. Burkart, A., Imhoff, A. B. & Roscher, E., “Foreign-body reaction    to the bioabsorbable suretac device,” Arthroscopy 16, 91-95 (2000).-   22. Butler, D. L., Goldstein, S. A. & Guilak, F., “Functional tissue    engineering: the role of biomechanics,” J. Biomech. Eng. 122,    570-575 (2000).-   23. Chen, C. H. et al., “Enveloping the tendon graft with periosteum    to enhance tendon-bone healing in a bone tunnel: A biomechanical and    histologic study in rabbits,” Arthroscopy 19, 290-296 (2003).-   24. Cheung, H. S. & McCarty, D. J., “Mitogenesis induced by    calcium-containing crystals. Role of intracellular dissolution,”    Exp. Cell Res. 157, 63-70 (1985).-   25. Clark, J. M. & Sidles, J. A., “The interrelation of fiber    bundles in the anterior cruciate ligament,” J. Orthop. Res. 8,    180-188 (1990).-   26. Cooper, J. A., “Design, optimization and in vivo evaluation of a    tissue-engineered anterior cruciate ligament replacement,” Drexel    University (Thesis/Dissertation) (2002).-   27. Cooper, J. A., Lu, H. H. & Laurencin, C. T., “Fiber-based tissue    engineering scaffold for ligament replacement: design considerations    and in vitro evaluation,” Proceedings of 5th World Biomaterial    Congress, 208 (Abstract)(2000).-   28. Daniel, D. M. et al., “Fate of the ACL-injured patient. A    prospective outcome study,” Am. J. Sports Med. 22, 632-644 (1994).-   29. Deitzel et al., Polymer 42, (2001).-   30. Deitzel et al., Polymer 43, (2002).-   31. Ducheyne, P., “Bioceramics: material characteristics versus in    vivo behavior,” J. Biomed. Matls. Res. 21, 219-236 (1987).-   32. Dunn, M. G., Liesch, J. B., Tiku, M. L., Maxian, S. H. &    Zawadsky, J. P., “The Tissue Engineering Approach to Ligament    Reconstruction,” Matls. Res. Soc. 331, 13-18 (1994).-   33. El-Amin, S. F. et al., “Human osteoblast integrin expression on    degradable polymeric materials for tissue engineered bone,” J.    Orthop. Res. (In Press)(2001).-   34. Erben, J. Histochem. Cytochem. 45, 307-314 (1997).-   35. Fleming, B. C., Abate, J. A., Peura, G. D. & Beynnon, B. D.,    “The relationship between graft tensioning and the    anterior-posterior laxity in the anterior cruciate ligament    reconstructed goat knee,” J. Orthop. Res. 19, 841-844 (2001).-   36. Fleming, B., Beynnon, B., Howe, J., McLeod, W. & Pope, M.,    “Effect of tension and placement of a prosthetic anterior cruciate    ligament on the anteroposterior laxity of the knee,” J. Orthop. Res.    10, 177-186 (1992).-   37. Fong et al., Polymer, (1999).-   38. Fridrikh et al., PhysRevLet (2003).-   39. Fu, F. H., Bennett, C. H., Ma, C. B., Menetrey, J. & Lattermann,    C., “Current trends in anterior cruciate ligament reconstruction.    Part II. Operative procedures and clinical correlations,” Am. J.    Sports Med. 28, 124 130 (2000).-   40. Fujikawa, K., Iseki, F. & Seedhom, B. B., “Arthroscopy after    anterior cruciate reconstruction with the Leeds-Keio ligament,” J.    Bone Joint Surg. Br. 71, 566-570 (1989).-   41. Gao, J. & Messner, K., “Quantitative comparison of soft    tissue-bone interface at chondral ligament insertions in the rabbit    knee joint,” J. Anat. 188, 367-373 (1996).-   42. Gao, J., Rasanen, T., Persliden, J. & Messner, K., “The    morphology of ligament insertions after failure at low strain    velocity: an evaluation of ligament entheses in the rabbit knee,” J.    Anat. 189, 127-133 (1996).-   43. Gotlin, R. S. & Huie, G., “Anterior cruciate ligament injuries.    Operative and rehabilitative options,” Phys. Med. Rehabil. Clin. N.    Am. 11, 895-928 (2000).-   44. Goulet, F. et al., “Principles of Tissue Engineering,” Lanza, R.    P., Langer, R. & Vacanti, J. P. (eds.), pp. 639-645 Academic Press    (2000).-   45. Grana, W. A., Egle, D. M., Mahnken, R. & Goodhart, C. W., “An    analysis of autograft fixation after anterior cruciate ligament    reconstruction in a rabbit model,” Am. J. Sports Med. 22, 344-351    (1994).-   46. Gregoire, M., Orly, I., Kerebel, L. M. & Kerebel, B., “In vitro    effects of calcium phosphate biomaterials on fibroblastic cell    behavior,” Biol. Cell 59, 255-260 (1987).-   47. Harner, C. D. et al., “Quantitative analysis of human cruciate    ligament insertions,” Arthroscopy 15, 741-749 (1999).-   48. Hench, L. L., “Bioceramics: from concept to clinic,” J. Am.    Cera. Soc. 74(7), 1487-1510 (1991).-   49. Jackson, D. W., American Academy of Orthopaedic Surgeon Bulletin    40, 10-11 (1992).-   50. Jackson, D. W., Grood, E. S., Arnoczky, S. P., Butler, D. L. &    Simon, T. M., “Cruciate reconstruction using freeze dried anterior    cruciate ligament allograft and a ligament augmentation device    (LAD). An experimental study in a goat model,” Am. J. Sports Med.    15, 528-538 (1987).-   51. Jackson, D. W. et al., Trans. Orhtop. Res. Soc. 16, 208    (Abstract) (1991).-   52. Jackson, D. W. et al., “A comparison of patellar tendon    autograft and allograft used for anterior cruciate ligament    reconstruction in the goat model,” Am. J. Sports Med. 21, 176-185    (1993).-   53. Jiang, J., Nicoll, S. B. & Lu, H. H., “Effects of Osteoblast and    Chondrocyte Co-Culture on Chondrogenic and Osteoblastic Phenotype In    Vitro,” Trans. Orhtop. Res. Soc. 49 (Abstract) (2003).-   54. Johnson, R. J., “The anterior cruciate: a dilemma in sports    medicine,” Int. J. Sports Med. 3, 71-79 (1982).-   55. Kim et al., Biomaterials (2003).-   56. Kurosaka, M., Yoshiya, S. & Andrish, J. T., “A biomechanical    comparison of different surgical techniques of graft fixation in    anterior cruciate ligament reconstruction,” Am. J. Sports Med. 15,    225-229 (1987).-   57. Kurzweil, P. R., Frogameni, A. D. & Jackson, D. W., “Tibial    interference screw removal following anterior cruciate ligament    reconstruction,” Arthroscopy 11, 289-291 (1995).-   58. Larson, R. P., “The Crucial Ligaments: Diagnosis and Treatment    of Ligamentous Injuries About the Knee,”John, A. Jr. & Feagin, J. A.    (eds.), pp. 785-796 Churchill Livingstone, N.Y. (1994).-   59. Liu, S. H. et al., “Morphology and matrix composition during    early tendon to bone healing,” Clinical Orthopaedics & Related.    Research. 253-260 (1997).-   60. Loh, J. C. et al., “Knee stability and graft function following    anterior cruciate ligament reconstruction: Comparison between 11    o'clock and 10 o'clock femoral tunnel placement,” Arthroscopy 19,    297-304 (2003).-   61. Lu, H. H., Pollack, S. R. & Ducheyne, P., “Temporal zeta    potential variations of 45S5 bioactive glass immersed in an    electrolyte solution,” J. Biomed. Matls. Res. 51, 80-87 (2000).-   62. Lu, H. H., Pollack, S. R. & Ducheyne, P., “45S5 bioactive glass    surface charge variations and the formation of a surface calcium    phosphate layer in a solution containing fibronectin,” J. Biomed.    Matls. Res. 54, 454-461 (2001).-   63. Lu, H. H., Cooper, J. A., Ko, F. A., Attawia, M. A. &    Laurencin, C. T., “Effect of polymer scaffold composition on the    morphology and growth of anterior cruciate ligaments cells,” Society    for Biomaterials Proceedings (Abstract) (2001).-   64. Lu, H. H., El Amin, S. F., Scott, K. D. & Laurencin, C. T.,    “Three dimensional, bioactive, biodegradable, polymer bioactive    glass composite scaffolds with improved mechanical properties    support collagen synthesis and mineralization of human    osteoblast-like cells in vitro,” J. Biomed. Matls. Res. 64A, 465-474    (2003).-   65. Lu, H. H. et al., “Evaluation of Optimal Parameters in the    Co-Culture of Human Anterior Cruciate Ligament Fibroblasts and    Osteoblasts for Interface Tissue Engineering,” ASME 2003 Summer    Bioengineering Conference (Abstract) (2003).-   66. Markolf, K. L. et al., “Effects of femoral tunnel placement on    knee laxity and forces in an anterior cruciate ligament graft,” J.    Orthop. Res. 20, 1016-1024 (2002).-   67. Matthews, L. S., Soffer, S. R., “Pitfalls in the use of    interference screws for anterior cruciate ligament reconstruction:    brief report,” Arthroscopy 5, 225-226 (1989).-   68. Matyas, J. R., Anton, M. G., Shrive, N. G. & Frank, C. B.,    “Stress governs tissue phenotype at the femoral insertion of the    rabbit MCL,” J. Biomech. 28, 147-157 (1995).-   69. McCarthy, D. M., Tolin, B. S., Schwendeman, L., Friedman, M. J.    & Woo, S. L., “The Anterior Cruciate Ligament: Current and Future    Concepts,” Douglas, W. & MD Jackson (eds.) Raven Press, New York    (1993).-   70. Messner, K., “Postnatal development of the cruciate ligament    insertions in the rat knee. Morphological evaluation and    immunohistochemical study of collagens types I and II,” Acta    Anatomica 160, 261-268 (1997).-   71. Moore, P. B. & Dedman, J. R., “Calcium binding proteins and    cellular regulation,” Life Sci. 31, 2937-2946 (1982).-   72. Nicoll, S. B., Wedrychowska, A., Smith, N. R. & Bhatnagar, R.    S., “Modulation of proteoglycan and collagen profiles in human    dermal fibroblasts by high density micromass culture and treatment    with lactic acid suggests change to a chondrogenic phenotype,”    Connect. Tissue Res. 42, 59-69 (2001).-   73. Niyibizi, C., Sagarrigo, V. C., Gibson, G. & Kavalkovich, K.,    “Identification and immunolocalization of type X collagen at the    ligament-bone interface,” Biochem. Biophys. Res. Commun. 222,    584-589 (1996).-   74. Noyes, F. R. & Barber-Westin, S. D., “Revision anterior cruciate    ligament surgery: experience from Cincinnati,” Clin. Orthop. 116-129    (1996).-   75. Noyes, F. R., Mangine, R. E. & Barber, S., “Early knee motion    after open and arthroscopic anterior cruciate ligament    reconstruction,” Am. J. Sports Med. 15, 149-160 (1987).-   76. Panni, A. S., Milano, G., Lucania, L. & Fabbriciani, C., “Graft    healing after anterior cruciate ligament reconstruction in rabbits,”    Clin. Orthop. 203-212 (1997).-   77. Pena, F., Grontvedt, T., Brown, G. A., Aune, A. K. &    Engebretsen, L., “Comparison of failure strength between metallic    and absorbable interference screws. Influence of insertion torque,    tunnel-bone block gap, bone mineral density, and interference,”    Am. J. Sports Med. 24, 329-334 (1996).-   78. Petersen, W. & Tillmann, B., “Structure and vascularization of    the cruciate ligaments of the human knee joint,” Anat. Embryol.    (Berl) 200, 325-334 (1999).-   79. Robertson, D. B., Daniel, D. M. & Biden, E., “Soft tissue    fixation to bone,” Am. J. Sports Med. 14, 398-403 (1986).-   80. Rodeo, S. A., Suzuki, K., Deng, X. H., Wozney, J. & Warren, R.    F., “Use of recombinant human bone morphogenetic protein-2 to    enhance tendon healing in a bone tunnel,” Am. J. Sports Med. 27,    476-488 (1999).-   81. Rodeo, S. A., Arnoczky, S. P., Torzilli, P. A., Hidaka, C. &    Warren, R. F., “Tendon-healing in a bone tunnel. A biomechanical and    histological study in the dog,” J. Bone Joint Surg. Am. 75,    1795-1803 (1993).-   82. Safran, M. A. & Harner, C. D., “Technical considerations of    revision anterior cruciate ligament surgery,” Clin. Orthop. 50-64    (1996).-   83. Sagarriga, V. C., Kavalkovich, K., Wu, J. & Niyibizi, C.,    “Biochemical analysis of collagens at the ligament-bone interface    reveals presence of cartilage-specific collagens,” Arch. Biochem.    Biophys. 328, 135-142 (1996).-   84. Scapinelli, R. & Little, K., “Observations on the mechanically    induced differentiation of cartilage from fibrous connective    tissue,” J. Pathol. 101, 85-91 (1970).-   85. Schafer, et al., “In vitro generation of osteochondral    composites,” Biomaterials 21:2599-2606 (2000).-   86. Schafer, et al., “Tissue-engineered composites for the repair of    large osteochondral defects,” Arthritis Rheum. 46:2524-2534 (2002).-   87. Shellock, F. G., Mink, J. H., Curtin, S. & Friedman, M. J., “MR    imaging and metallic implants for anterior cruciate ligament    reconstruction: assessment of ferromagnetism and artifact,” J. Magn.    Reson. Imaging 2, 225-228 (1992).-   88. Shin et al., Polymer 42, (2001).-   89. Sittinger, et al., “Engineering of cartilage tissue using    bieresorbable polymer carriers in perusion culture,” Biomaterials    15(6):451-456 (1994).-   90. Spalazzi, J. P., Dionisio, K. L., Jiang, J. & Lu, H. H.,    “Chondrocyte and Osteoblast Interaction on a Degradable Polymer    Ceramic Scaffold,” ASME 2003 Summer Bioengineering Conference    (Abstract) (2003).-   91. Steiner, M. E., Hecker, A. T., Brown, C. H., Jr. & Hayes, W. C.,    “Anterior cruciate ligament graft fixation. Comparison of hamstring    and patellar tendon grafts,” Am. J. Sports Med. 22, 240-246 (1994).-   92. Thomas, N. P., Turner, T. G. & Jones, C. B., “Prosthetic    anterior cruciate ligaments in the rabbit. A comparison of four    types of replacement,” J. Bone Joint Surg. Br. 69, 312-316 (1987).-   93. Thomopoulos, S. et al., “The localized expression of    extracellular matrix components in healing tendon insertion sites:    an in situ hybridization study,” J. Orthop. Res. 20, 454-463 (2002).-   94. Wei, X. & Messner, K., “The postnatal development of the    insertions of the medial collateral ligament in the rat knee,” Anat.    Embryol. (Berl) 193, 53-59 (1996).-   95. Weiler, A., Hoffmann, R. F., Bail, H. J., Rehm, O. & Sudkamp, N.    P., “Tendon healing in a bone tunnel. Part II: Histologic analysis    after biodegradable interference fit fixation in a model of anterior    cruciate ligament reconstruction in sheep,” Arthroscopy 18, 124-135    (2002).-   96. Weiler, A., Windhagen, H. J., Raschke, M. J., Laumeyer, A. &    Hoffmann, R. F., “Biodegradable interference screw fixation exhibits    pull-out force and stiffness similar to titanium screws,” Am. J.    Sports Med. 26, 119-126 (1998).-   97. Woo, S. L., Gomez, M. A., Seguchi, Y., Endo, C. M. & Akeson, W.    H., “Measurement of mechanical properties of ligament substance from    a bone-ligament-bone preparation,” J. Orthop. Res. 1, 22-29 (1983).-   98. Woo, S. L., Newton, P. O., MacKenna, D. A. & Lyon, R. M., “A    comparative evaluation of the mechanical properties of the rabbit    medial collateral and anterior cruciate ligaments,” J. Biomech. 25,    377-386 (1992).-   99. Wu H. et al., J. Am. Chem. Soc. 2003 Jan. 15; 125(2):554-559.-   100. Wuthier, R. E., “Involvement of cellular metabolism of calcium    and phosphate in calcification of avian growth plate cartilage,” J.    Nutr. 123, 301-309 (1993).-   101. Xu et al., Biomaterials 25, 877-886, (2004).-   102. Yahia, L., “Ligaments and Ligamentoplasties,” Springer Verlag,    Berlin Heidelberg (1997).-   103. Yoshiya, S., Nagano, M., Kurosaka, M., Muratsu, H. & Mizuno,    K., “Graft healing in the bone tunnel in anterior cruciate ligament    reconstruction,” Clin. Orthop. 278-286 (2000).

1-89. (canceled)
 90. A scaffold apparatus for musculoskeletal tissue engineering comprising a plurality of phases, wherein the apparatus has a gradient of mineral content across the phases, and at least one of the phases comprises a polymer-ceramic composite material.
 91. The scaffold apparatus of claim 90, wherein the mineral content is calcium phosphate content.
 92. The scaffold apparatus of claim 90, integrated in a graft fixation device.
 93. The scaffold apparatus of claim 90, integrated in a graft collar. 